Hydroxyphenyl cross-linked macromolecular network and applications thereof

ABSTRACT

A dihydroxyphenyl cross-linked macromolecular network is provided that is useful in artificial tissue and tissue engineering applications, particularly to provide a synthetic macromolecular network for a wide variety of tissue types. In particular, artificial or synthetic cartilage, vocal cord material, vitreous material, soft tissue material and mitral valve material are described. In an embodiment, the network is composed of tyramine-substituted and cross-linked hyaluronan molecules, wherein cross-linking is achieved via peroxidase-mediated dityramine-linkages that can be performed in vivo. The dityramine bonds provide a stable, coherent hyaluronan-based hydrogel with desired physical properties.

This application is a continuation of co-pending application Ser. No.12/283,661 filed Sep. 15, 2008, which is a continuation of applicationSer. No. 11/176,544 filed Jul. 7, 2005 (now U.S. Pat. No. 7,465,665),which is a continuation-in-part of application Ser. No. 10/753,779 filedJan. 8, 2004 (now U.S. Pat. No. 6,982,298), which application claims thebenefit of U.S. Provisional Patent Application Ser. No. 60/439,201 filedon Jan. 10, 2003.

BACKGROUND OF THE INVENTION

Articular cartilage performs an essential function in healthy joints. Itis responsible for absorbing and dissipating impact and frictional loadsin order to divert these loads away from bones, to protect the bonesfrom damage. Cartilage performs this function by transferring theloading force to a fluid phase within a three-dimensional network ofaggrecan molecules, themselves constrained (described in the nextparagraph) within the joint space. Aggrecan molecules have up to 100chondroitin sulfate (CS) chains attached to a core protein, with eachchondroitin sulfate chain possessing multiple negatively charged sulfategroups along their length. The effect of all these sulfate groups is tocause each of the chondroitin sulfate chains in a single aggrecanmolecule to repel one another, (resulting in the aggrecan moleculehaving the maximum possible volume at rest), and also to cause adjacentaggrecan molecules in a cartilage aggregate to repel one another.

In healthy cartilage, aggrecan molecules are attached to long hyaluronanchains, which are in turn constrained in large cartilage aggregateswithin the joint space by an extracellular collagen fibril matrix. Thus,even though adjacent chondroitin sulfate chains in each aggrecanmolecule (and adjacent aggrecan molecules attached to the same or adifferent hyaluronan chain) repel one another, they are nonethelessconstrained within the collagen matrix. See FIG. 1 depicting normal,healthy cartilage. Because the chondroitin sulfate chains are sorepulsive, the hyaluronan-aggrecan network (or macromolecular network)expands as much as possible within the constraints of the collagenmatrix to achieve the lowest possible energy state at rest; i.e. toallow the maximum possible spacing between adjacent negatively chargedsulfate groups. As a result, network molecules are highly resistant tobeing shifted or displaced in order to avoid approaching an adjacentnetwork molecule. These large cartilage aggregates are trapped at onefifth their free solution volume within a meshwork of collagen fibers,which resist any further swelling. Cartilage aggregates with their highnegative charge density bind large solvent domains, and contribute tocartilage's ability to absorb loads and resist deformation. Uponcompression, the distance between the fixed-negative charge groups onthe proteoglycans decreases, which increases the charge-to-chargerepulsive forces as well as the concentration of free-floating positivecounterions (such as Ca²⁺ and Na⁺). Both effects contribute to theviscoelastic nature of cartilage and its ability to resist deformationand absorb compressive loads, further described below.

Within the macromolecular network are water molecules which provide asubstantially continuous fluid phase. The macromolecular network divertsimpact and frictional loads away from bones by transferring them to thecontinuous fluid (water) phase as follows. As a joint undergoes a load,the force is absorbed first by the macromolecular network, where it actson and tends to deform or compress the network. The force sets uppressure gradients in the fluid phase in order to induce fluid flow toaccommodate network deformation or compression resulting from the load.But the fluid cannot negotiate the tight macromolecular network, packedwith the repulsive chondroitin sulfate chains, sufficiently toaccommodate a bulk flow of water without shifting or displacing thenetwork molecules. Hence, individual water molecules may diffuse withinthe network, but the bulk fluid phase is substantially constrained fromflowing through the network except at a much slowed rate due to theresistance to displacement of network molecules. Because the watermolecules cannot flow readily despite the pressure gradients, the energyfrom the impact or frictional load is transferred to and absorbed by thefluid phase where it contributes to compressing the liquid water untilthe water can be sufficiently displaced to accommodate the networkconformation and the pressure gradients have subsided. The overallresult is that cartilage absorbs the potentially harmful load, therebydiverting it from bone.

Through this elegant mechanism, normal cartilage is capable of absorbingsignificant loads by transferring the bulk of the loading force to afluid phase constrained within a macromolecular network. Thisarrangement has yet to be adequately duplicated via artificial orsynthetic means in the prior art. Consequently, there is no adequateremedy for cartilage degenerative disorders, such as arthriticdisorders, where the aggrecan molecules become separated from theirhyaluronan chains and are digested or otherwise carried out from thecartilage aggregates.

Osteoarthritis and rheumatoid arthritis affect an estimated 20.7 and 2.1million Americans, respectively. Osteoarthritis alone is responsible forroughly 7 million physician visits a year. For severe disablingarthritis, current treatment involves total joint replacement with onaverage 168,000 total hip replacements and 267,000 total kneereplacements performed per year in the U.S. alone. Defects in articularcartilage present a complicated treatment problem because of the limitedcapacity of chondrocytes to repair cartilage. Treatment strategies todate have focused on the use of autologous chondrocytes expanded inculture or the recruitment of mesenchymal stem cells in vivo bychemotactic or mitogenic agents. The intent of these strategies is toincrease and/or activate the chondrocyte population so as toresynthesize a normal, healthy articular cartilage surface. One majordifficulty associated with these strategies is the inability to maintainthese agents at the site of the defect. Hyaluronan has been proposed asa candidate for the development of biomaterials for local delivery ofchondrocytes or bioactive agents because of its unique properties,including excellent biocompatibility, degradability, and rheological andphysiochemical properties. However, it has been unknown whetherchondrocytes suspended in a tissue engineered hyaluronan matrix would beable to synthesize a new cartilage matrix with mechanical propertiescomparable to normal, healthy articular cartilage. This is becauseconventional biomaterials made from hyaluronan are formed throughchemistries that are incompatible with maintaining cell viability.Chondrocytes must be introduced to the matrices after matrix formationwith variable and normally poor results.

Accordingly, there is a need in the art for an artificial or syntheticmatrix that can effectively divert a loading force from bones in aneffective manner. Preferably, such a matrix can be provided in situ orin vivo to repair or replace articular cartilage during an orthopedicsurgical procedure. Most preferably, the artificial or synthetic matrixcan be provided to an in situ or in vivo target site as a liquid or aplurality of liquids, and can set up in place to provide a substantiallyseamless integration with existing cartilaginous and/or bony tissue in apatient.

It also is desirable to provide an artificial or synthetic matrix thatcan be used or adapted to synthesize a variety of replacement tissues.

SUMMARY OF THE INVENTION

A synthetic macromolecular network is provided including amacromolecular network that includes the following structure

wherein R₁ and R₂ each comprises a structure selected from the groupconsisting of polycarboxylates, polyamines, polyhydroxyphenyl molecules,and copolymers thereof, and wherein R₁ and R₂ can be the same ordifferent structures.

More generally, a synthetic macromolecular network is provided thatcomprises macromolecules selected from the group consisting ofhydroxyphenyl-substituted polycarboxylates, hydroxyphenyl-substitutedpolyamines, other polyhydroxyphenyl molecules, and copolymers thereof,as well as at least one dihydroxyphenyl linkage linking differentmacromolecules or linking different sites on the same macromolecule.

A variety of synthetic, implantable tissue materials also are providedwhich include or are composed of the macromolecular network mentioned inthe preceding paragraph, including a synthetic, implantable cartilagematerial; a synthetic, implantable vocal cord material; a synthetic,implantable vitreous material; a synthetic, implantable soft tissuematerial; and a synthetic, implantable mitral valve material.

BRIEF DESCRIPTION OF THE DRAWINGS

The file of this patent contains at least one drawing executed in color.Copies of this patent with color drawings will be provided by the PatentOffice upon request and payment of the necessary fee.

FIG. 1 is a schematic diagram of normal, healthy human cartilage.

FIG. 2 is a schematic diagram of a dihydroxyphenyl cross-linkedmacromolecular network according to the invention.

FIG. 3 is a structural formula of a hyaluronan molecule.

FIGS. 4 a-4 c are graphs showing comparative results for mechanicaltesting in a confined compression test (equilibrium stress versusapplied strain) of T-HA (FIG. 4 a), T-Aggrecan (FIG. 4 b) and 50%T-HA/50% T-Aggrecan composite (FIG. 4 c) hydrogels according to theinvention versus published results for articular cartilage plugs(Example 3). The relationship between glycosaminoglycan (GAG)concentration and material compressive strength is shown in FIG. 4 d.

FIG. 5 is a graph showing comparative data of glucose utilization forchondrocytes embedded in T-HA hydrogels (1.7% and 4.7% T-HA) compared tocultured on tissue culture plastic (control).

FIG. 6 is a series of four photographs illustrating a surgical procedureto implant a T-HA hydrogel into articular cartilage defects according toan aspect of the invention described in Example 6.

FIG. 7 is a series of two photographs showing the T-HA hydrogel implantsone month after implantation into the medial trochlar facet of a Yucatanpig as described in Example 6, as well as the opposing (articulating)patella surface.

FIG. 8 is a series of photographs illustrating the histological resultsof control side (unfilled) and experimental side (TB-HA hydrogel filled)canine vocal cords, 3 months post-operatively, following a vocal cordrepair procedure using a T-HA hydrogel as a synthetic vocal cordmaterial as described in Example 7.

FIG. 9 is a series of photographs illustrating the histological resultsof surgically augmented vocal cords in a rabbit model using a T-HAhydrogel as a synthetic vocal cord material, also as described inExample 7.

FIG. 10 is a series of photographs of control (unoperated) andexperimental (surgically replaced) eyes one month post-operative,following a vitreous replacement procedure using T-HA hydrogel as asynthetic vitreous material as described in Example 8.

FIG. 11 shows comparative electroretinogram (ERG) results recorded forboth control and vitreous replaced eyes in response to flashes of lightin a rabbit model as described in Example 8.

FIG. 12 is a series of electron micrographs of the retina from fourquadrants of control (unoperated) and experimental (surgically replaced)eyes one month post-operative, following a vitreous replacementprocedure using T-HA hydrogel as a synthetic vitreous material asdescribed in Example 8.

FIG. 13 is a series of photographs showing representative results ofhistological results for a 100 mg/ml T-HA hydrogel plug implantedsubcutaneously into an immunocompetent rat at one month post-operativelyas described in Example 9.

FIG. 14 is a photograph of a cadaveric canine heart used to specify T-HAhydrogel materials for mitral valve repair as described in Example 10.

DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS OF THE INVENTION

As used herein, the term polycarboxylate means a molecule, structure orspecies having a chain length of at least two functional groups orunits, wherein at least two such groups or units of the chain are orcomprise carboxylic acid groups that are sterically accessible to anucleophilic substitution reaction as described herein. Also as usedherein, the term polyamine means a molecule, structure or species havinga chain length of at least two functional groups or units, wherein atleast two such groups or units of the chain are or comprise primaryamine groups that are available for a nucleophilic substitutionreaction. Also as used herein, a polyhydroxyphenyl molecule means amolecule having a chain length of at least two functional groups orunits, wherein at least two such groups or units of the chain are orcomprise hydroxyphenyl groups that can be linked to anotherhydroxyphenyl group via a C—C bond. Also as used herein, a hydrogel is amaterial that is prepared comprising a macromolecular network that isused or useful in tissue replacement or engineering applications, e.g.as artificial cartilage, as a material to coat surgical instruments toprevent tissue irritation, or to provide a semi-permeable membrane suchas for use in an artificial kidney, etc.

The invention includes a novel structure of a macromolecular networkthat has been formed by linking hydroxyphenyl groups attached toadjacent long chain macromolecules, resulting in effectivelycross-linking the macromolecules to provide a large network. The basiccross-linking structure of the network is shown below

where R₁ and R₂ are each long chain macromolecules. R₁ and R₂ can be thesame molecule or different molecules, but it will be understood that toprovide a suitable network, R₁ and R₂ will be different molecules for atleast a portion of the dihydroxyphenyl linkages in a network accordingto the invention. It is not necessary, though it is preferred, that R₁and R₂ are the same species of molecule.

By providing a plurality of these dihydroxyphenyl linkages betweenadjacent macromolecules, a network of dihydroxyphenyl cross-linkedmacromolecules is provided as shown schematically in FIG. 2. In thefigure, the macromolecules are represented schematically by cylindricalstrands, each preferably having at least two hydroxyphenyl groupsattached along its length. It is noted that not every hydroxyphenylgroup must be linked to another hydroxyphenyl group.

Briefly, the disclosed invention involves covalent coupling ofhydroxyphenyl containing compounds, including but not limited totyramine, through their primary amine (or carboxyl) groups to carboxyl(or primary amine) groups on various polymeric scaffold materials,including but not limited to hyaluronan or chondroitin sulfate (e.g. inthe form of aggrecan), via a carbodiimide-mediated reaction. Afterisolation and purification of the hydroxyphenyl-substituted polymericscaffolds, the hydroxyphenyl residues are selectively cross-linked byhorseradish peroxidase (HRP) in the presence of very dilute hydrogenperoxide to form hydrogels. As will become apparent, the hydrogels madeas described herein are or can be used as a fully implantable,non-immunogenic synthetic tissue matrix material that can be implantedinto the body for a variety of purposes as will be described. As usedherein, ‘implantable’ refers both to surgical implantation of a hydrogelas through a surgical incision, and to provision of the hydrogel withinthe body via injection, e.g. using a syringe. Whether surgicallyimplanted or injected, the implantable hydrogels can be provided withinthe body already cross-linked (ex vivo cross-linking) or otherwise itcan be cross-linked in situ at the site of implantation within the bodyas will be further described.

The first step in providing the macromolecular network is to prepare orprovide the long-chain macromolecules having periodic hydroxyphenylgroups attached. In one embodiment, the macromolecules arepolyhydroxyphenyl molecules which already have multiple or periodichydroxyphenyl groups, such as polyphenols. Suitable polyphenols includepolyamino acids (e.g. polytyrosine), epigallocatechin (EGC), andepigallocatechin gallate (EGCG) isolated from green tea, less preferablyother polyphenols.

In a further embodiment, the hydroxyphenyl groups can be added to themacromolecules periodically or randomly along their length via achemical reaction. A preferred method of adding hydroxyphenyl groups tothe macromolecules is to utilize a carbodiimide-mediated substitutionreaction pathway to provide an amide bond between a primary amine havinga hydroxyphenyl group and a carboxylic acid group attached to themacromolecules. In this method, the long-chain macromolecule preferablyis a polycarboxylate molecule, having periodic carboxylic acid groupsalong its length. The hydroxyphenyl groups are provided as part ofsmaller molecules having primary amine groups that can be attached tothe carboxyl carbon atoms of a carboxylic acid group on the long-chainmacromolecules via the carbodiimide pathway. The reaction proceeds asfollows:

where:

Structure A is a carbodiimide;

Structure B is a polycarboxylate (though only one CO₂H group is shown);

Structure C is the product of Reaction A and is an activatedO-acylisourea;

Structure D is a primary amine having a hydroxyphenyl group;

Structure E is a hydroxyphenyl-substituted polycarboxylate; and

Structure F is an acylurea byproduct;

wherein individual Rs can be individually selected, the same ordifferent from one another, to be a straight chain or branched alkane oracyl group, or any other structure that does not interfere with thecarbodiimide reaction pathway to provide the amide bond between the NH₂and CO₂H groups as shown in Structure E above.

In the above-illustrated pathway, Reaction A represents a carbodiimideactivation of the carboxyl group to provide an activated O-acylisoureaintermediate. The electropositive carbon atom of this intermediate isreceptive to nucleophilic attack by the lone pair of electrons on anitrogen atom of an adjacent primary amine molecule having an attachedhydroxyphenyl group. The products of this nucleophilic substitutionreaction (Reaction B) are a hydroxyphenyl-substituted polycarboxylateand an acylurea byproduct which can be dialyzed out to provide asubstantially pure hydroxyphenyl-substituted polycarboxylate product.

Certain side-reactions are possible in the above-described carbodiimidereaction pathway chemistry and should be considered by the person havingordinary skill in the art. First, the carbodiimide can react withnucleophiles other than the carboxylate oxygen atom of thepolycarboxylate molecule required to form the desired O-acylisourea(reaction A). Such nucleophiles may include the amine and/orhydroxyphenyl groups of Structure D illustrated above. In particular,there are three potential side-reactions for Reaction A which can reducethe effective concentration of the carbodiimide and the primary aminehaving the hydroxyphenyl group (Structures A and D), and potentiallylead to the creation of undesired adducts on the polycarboxylate(Structure B):

The product of an amine reaction with the carbodiimide (Reaction C) willnot have a free amine group effectively reducing the amount of tyramineavailable for reaction with the O-acylisourea. This reaction alsoreduces the amount of carbodiimide available for formation of thedesired O-acylisourea. The products of the hydroxyphenyl reaction(Reaction D) are not UV absorbent, which will make their detection byUV-spectroscopy in the final hydroxyphenyl-substituted polycarboxylateproduct (explained below) more difficult. However, because theseproducts still contain free amine groups, they can form amide bonds withthe polycarboxylate molecule via Reaction B. This can give rise to twounproductive hyaluronan-substituted structures, neither of which canparticipate in the peroxidase cross-linking reaction in the second step(described below) of preparing the cross-linked network due to theabsence of an extractable phenolic hydroxyl hydrogen atom needed togenerate the free radical (also explained below). Finally, thecarbodiimide can react non-productively with water (Reaction E) toproduce the same acylurea shown above as a byproduct of Reaction B, butwith none of Structure E, the desired product.

Once the desired O-acylisourea product has been formed in Reaction A,there is again the possibility for certain additional side-reactions:

The O-acylisourea (Structure C) can be hydrolyzed as shown in Reaction Freleasing the original unmodified polycarboxylate (Structure B) and theacylurea of the carbodiimide (Structure F). This is an unproductivereaction similar to reaction E, which reduces the effectiveconcentration of the carbodiimide. The O-acylisourea, can also undergoan intramolecular rearrangement (Reaction G) to form two unreactiveN-acylureas. These structures form unproductive adducts on thecarboxylate molecule which cannot contribute to the peroxidase catalyzedcross-linking reaction (step 2 discussed below) for preparing a networkaccording to the invention. The O-acylisourea can also react (ReactionH) with a second carboxyl group on either the same or a differentpolycarboxylate molecule to form an acid anhydride. This molecule canthen react with Structure D to form the desired amide and regenerate thesecond carboxyl group. Thus there are two potential side-reactions forthe O-acylisourea, which can reduce the effective concentration of thecarbodiimide (Reactions F and G), and potentially lead to creation ofundesired adducts on the polycarboxylate molecule.

Negative effects of these side reactions can be addressed throughconventional techniques without undue experimentation.

Alternatively to the pathway shown above where the macromolecule(Structure B) is a polycarboxylate, the macromolecule can be a polyaminehaving multiple or periodic amine groups along its length, wherein thehydroxyphenyl groups then are provided as part of smaller carboxylicacid molecules. Suitable polyamines include: polyhexosamines such aschitosan (polyglucosamine); polyamino acids such as polylysine;polydeoxyribonucleotides such as poly (dA) (polydeoxyadenylic acid),poly(dC) (polydeoxycytidylic acid), and poly(dG) (polydeoxyguanylicacid); and polyribonucleotides such as poly(A) (polyadenylic acid),poly(C) (polycytidylic acid), and poly(G) (polyguanylic acid). Thecarbodimide-mediated reaction pathway proceeds exactly as explainedabove to form the amide bond between the amine group and carboxylic acidgroup except that, as will be understood by a person having ordinaryskill in the art, the resulting product will behydroxyphenyl-substituted polyamine instead of a polycarboxylate. Otherpeptides and/or proteins also can be used as the macromolecules in thepresent invention, either which have hydroxyphenyl groups disposed alongtheir length, or to which hydroxyphenyl groups can be provided via asubstitution reaction as described herein. For example, in addition tothe peptides already disclosed herein, polyarginine can be used as themacromolecule.

When substituting onto a polycarboxylate molecule, suitablehydroxyphenyl-containing compounds for use in the present inventioninclude those having a free primary amine that can be used to modifyscaffold materials having multiple or periodic CO₂H groups, includingtyrosine (2-amino-3-(4-hydroxyphenyl) proprionic acid) and tyramine(tyrosamine or 2-(4-hydroxyphenyl)ethylamine). When substituting onto apolyamine, suitable hydroxyphenyl-containing compounds include thosehaving a free CO₂H group that can be used to modify scaffold materialshaving multiple or periodic primary NH₂ groups, including tyrosine,3-(4-hydroxyphenyl) propionic acid and 4-hydroxyphenylacetic acid.

The second step in preparing a cross-linked macromolecular networkaccording to the invention is to link the resulting macromolecules, nowhaving one or more hydroxyphenyl groups attached, via a dihydroxyphenyllinking structure. In this step hydroxyphenyl groups attached todifferent macromolecules are linked via the reaction mechanism shownbelow using a peroxide reagent in the presence of a peroxidase:

(It is noted that some dihydroxyphenyl linking may occur betweendifferent hydroxyphenyl groups attached to the same molecule as well).Peroxidase in the presence of a dilute peroxide (preferably H₂O₂) isable to extract the phenolic hydroxyl hydrogen atom from hydroxyphenylcontaining compounds (such as tyramine) leaving the phenolic hydroxyloxygen with a single unshared electron, an extremely reactive freeradical. The free radical isomerizes to one of the two equivalentortho-position carbons and then two such structures dimerize to form acovalent bond effectively cross-linking the structures, which afterenolizing generates a dihydroxyphenyl dimer (a dihydroxyphenyl linkagesuch as dityramine linkage as described below).

For clarity, only a single dihydroxyphenyl linking reaction is shownabove, but it will be understood that several or multiple such linkageswill be produced when macromolecules having attached hydroxyphenylgroups are subjected to the reaction conditions (peroxide andperoxidase). Hydrogen peroxide is indicated in the above mechanism, butother suitable peroxides can be used. Also, the peroxidase preferably ishorseradish peroxidase (HRP). Alternatively, any other suitable enzyme(or other agent) can be used that is capable of generating free-radicalsfor cross-linking scaffold materials containing hydroxyphenyl groups,preferably under ordinary metabolic conditions as described below.

We have shown that the interaction of horseradish peroxidase (Type II)and hydrogen peroxide (H₂O₂) is suitable for the production ofcross-linked macromolecular networks. The mechanism comprises fourdistinct steps: (a) binding of peroxide to the heme-Fe(III) complex ofthe peroxidase to form an unstable peroxide complex, “Compound I”; (b)oxidation of the iron to generate a ferryl species with a pi-cationradical in the heme porphyrin ring, “Compound II”; (c) reduction ofCompound II by one substrate (i.e. hydroxyphenyl or water) molecule toproduce a product (i.e. hydroxyphenyl or superoxide) radical and anotherferryl species, “Compound III”; (d) reduction of Compound III by asecond substrate (i.e. hydroxyphenyl or water) molecule to release asecond product (i.e. hydroxyphenyl or superoxide) radical and regeneratethe native enzyme. Thus the peroxidase enzyme can either formhydroxyphenyl radicals required for cross-linking through interaction ofhydroxyphenyl groups at the enzyme active site to directly create thedesired radicals or through first generation of superoxide radicals,which then diffuse from the enzyme and interact with hydroxyphenylgroups to generate the desired radicals. Other compounds that have thepotential to produce the same effect include any porphyrin containingcompound (i.e. Photofrin below), which includes the peroxidase family,hemoproteins, or the structurally related chlorine compounds.

A number of other free radical initiators can be used to crosslink thehydroxyphenyl modified macromolecules described herein. A majority arebased on the formation or inclusion of reactive oxygen species (ROS)such as, but not limited to, molecules of hydrogen peroxide, ions ofhypochlorite, radicals like the hydroxyl radical, and the superoxideanion which is both ion and radical. Additional reactive molecules suchas reactive nitrogen species or reactive sulfur species, or those freeradical species involved in synthetic polymerization have the potentialto be used for hydroxyphenyl cross-linking.

ROS are commonly produced in nature through the use of enzymes, andsubstrates. Additional enzymatic systems which have the potential to beused in the cross-linking process, as a result of production ofsuperoxide radicals, include, but are not limited to xanthine-xanthineoxidase and NADPH-NADPH oxidase.

Another class of ROS free radical initiators that can be used involvesthe use of metallic cations. One example is based on the Fentonreaction, which takes place between hydrogen peroxide and a bivalentcation, such as Fe²⁺. This process generates powerful free radicals whenthe catalyst reacts with hydrogen peroxide. The principal chemicalreaction associated with Fenton's reaction is shown below:

H₂O₂+Fe²⁺=>OH.+OH⁻+Fe³⁺

where, Fe²⁺=ferrous ion, Fe³⁺=ferric ion, OH.=hydroxyl radicals

In addition to the initiation reaction described above that produceshydroxyl radicals, the Fenton's process can also produce superoxideradicals and hydroperoxide anions by additional chain propagationreactions described below. The perhydroxyl radical is known to be aweaker reductant compared to superoxide radical and hydroperoxideanions.

H₂O₂+OH.=>HO₂.+H₂O

HO₂.=>H⁺+O₂.⁻

HO₂.+O₂.⁻=>HO₂ ⁻+O₂

where O₂.⁻=superoxide radical anion, HO₂ ⁻=hydroperoxide anion,HO₂.=perhydroxyl radical.

We have demonstrated the ability for this reaction to crosslink tyraminesubstituted hyaluronan in the laboratory using ferrous sulfate inconjunction with hydrogen peroxide. Compounds which include, but are notlimited to, bivalent cations of copper, chromium, vanadium and cobaltcan be used in a similar manner. It is to be noted that while thehydroxyl free radical can be used to form a dityramine crosslink, it hasalso been shown to cleave HA chains, and thus may ultimately beunsuitable for ideal hydrogel formation.

Additional molecules or methods which can generate ROS include:

-   -   rubidium or cesium ions in the presence of oxygen to form        superoxide radicals;    -   trivalent cations, which with hydrogen peroxide form free        radicals and bivalent cations as shown below, which can        subsequently follow the reactions involved in the Fenton        process.

Fe⁺³+H₂O₂=Fe⁺²+.OOH+H⁺

-   -   the cytotoxic and antitumor therapy Photofrin, which upon        illumination with laser light at a wavelength of 630 nm causes        propagation of a radical generating reaction that produces        superoxide and hydroxyl radicals. In the absence of light, but        the presence of hydrogen peroxide, the porphorin ring in        Photofrin should operate by the same reaction as for the        peroxidase enzyme above.    -   UV light and hydrogen peroxide to form hydroxyl and superoxide        free radicals.    -   the persulfate family in combination with TEMED.

As noted above, one alternative method for generating such free-radicalsis to use Photofrin as an alternative, non-enzymatic, light-activatedcross-linking agent to cross-link the macromolecular network describedherein, e.g. tyramine-substituted hyaluronan to form tyraminecross-linked hyaluronan hydrogels. Photofrin®, which is known in theart, generates free radicals which could initiate the cross-linkingreaction as described herein in a manner similar to the peroxidase-H₂O₂mechanism described above. Photofrin® is a porfimer sodium manufacturedin powder or cake form by Wyeth-Ayerst Lederle Parenterals, Inc.

The dihydroxyphenyl cross-linked macromolecular network is superior toconventional cartilage or other tissue replacement or substitutionmethods and products, at least with respect to the ability to carry outan in situ cross-linking procedure, because the preferred cross-linkingreaction is enzyme driven (peroxidase). This means the cross-linkingreaction is carried out under ordinary in vivo or metabolic conditionsof temperature such as 35-39° C. (e.g. about 37° C.), pH range of 6-7(e.g. about 6.5), reagents etc. (A peroxide, such as hydrogen peroxide,is the only required reagent for the cross-linking reaction). Inaddition, Photofrin already is used in in vivo applications, e.g.ablative treatment of Barrett's esophagus, and the iron-basedcross-linking mechanism also can be optimized for in vivo performance.Thus, the cross-linking reaction can be performed in vivo, to provide across-linked hydrogel at a surgical situs, such as an orthopedicsurgical situs, to promote maximum seamless integration between thehydrogel and native tissue such as bony and cartilaginous tissue.Integration of the new hydrogel scaffold with native cartilage matrixmay occur immediately as the hydroxyphenyl-substituted macromolecularscaffold quickly penetrates into the existing cartilage matrix prior tocross-linking, and cross-links not only with otherhydroxyphenyl-substituted macromolecular scaffold material butpotentially with tyrosine residues of resident proteins in the existingcartilage matrix. This would eliminate a typical problem found withpre-formed matrix plugs, which is their poor integration into the nativecartilage tissue. The ability to cross-link the hydrogel directly on thearticular surface eliminates the need to surgically enlarge a defect tofit a pre-cast plug, as is necessary for hydrogels whose chemistries aretoxic to or otherwise prohibit their formation inside the patient. Itshould be noted that most cartilage damage as a result of arthritispresents as a variable thinning of the articular surface, not holes ofdefined shape.

For the peroxidase mechanism, because the cross-linking reactionrequires both the peroxide and a peroxidase (preferably horseradishperoxidase), solutions containing all but one of these components can beprepared for convenient application to a surgical site. For example, asolution comprising a tyramine—(or other hydroxyphenyl containingspecies) substituted polycarboxylate (such as tyramine-substitutedhyaluronan, etc.) and the peroxidase can be prepared, with a secondsolution prepared containing the peroxide. Alternatively, the peroxideand the peroxidase can be swapped between the first and secondsolutions, the important thing being that the peroxide and peroxidaseare kept separate (i.e. in separate solutions) until the cross-linkingreaction is to be carried out. Then, the first solution is applied,(e.g. to an in vivo surgical situs), and the second solution is appliedor sprayed over the first, in vivo, to cause in situ cross-linking ofthe tyramine residues. The cross linking reaction occurs in vivo. Othercombinations will be evident from the present disclosure which arewithin the skill of a person of ordinary skill in the art.

Furthermore, because the cross-linking reaction occurs under ordinarymetabolic conditions, additional living cells, such as chondrocytes,progenitor cells, stem cells, etc., can be provided directly to a mediumcontaining the non-cross-linked hydroxyphenyl-substitutedpolycarboxylates or polyamines (or polyphenols), i.e. to the first orsecond solution from the preceding paragraph, wherein the cell-richmedium is applied with the macromolecules to the site in vivo, and themolecules are subsequently cross-linked via addition of peroxidase andperoxide. The result is a cross-linked macromolecular network containingthe desired cells dispersed within it. Such a cell-enriched network isnot possible in conventional tissue replacement matrices due to theharsh conditions of temperature and pH under which they are prepared.Further, as described below in Example 5, it has been demonstrated thatthe cells provided to the network as described above remain viable evenafter cross-linking of tyramine-substituted hyaluronan (also describedbelow).

In a preferred embodiment particularly suitable for preparing syntheticcartilage as well as other synthetic or artificial tissues, themacromolecule used to produce the network is hyaluronan or hyaluronicacid (HA), and the hydroxyphenyl group is supplied in the form oftyramine. Hyaluronan (HA) is a ubiquitous molecule, which is mostconcentrated in specialized tissues such as cartilage, vocal cords,vitreous, synovial fluid, umbilical cord, and dermis. In these tissues,its function is manifold, influencing tissue viscosity, shockabsorption, wound healing, and space filling. HA has been shown toinfluence many processes within the extracellular matrix (ECM) in nativetissues where it is present including matrix assembly, cellproliferation, cell migration and embryonic/tissue development.

HA is composed of repeating pairs of glucuronic acid (glcA) andN-acetylglucosamine (glcNAc) residues linked by a β1,3 glycosidic bondas shown in FIG. 3. The glucuronic acid residue is particularlypertinent to the production of a macromolecular network as describedherein as this sugar provides an available carboxyl group periodicallyalong the repeat disaccharide structure of HA that is useful forhydroxyphenyl, i.e. tyramine, substitution. For each hyaluronan chain,this simple disaccharide is repeated up to 10,000 times or greaterresulting in macromolecule that can have a molecular weight on the order10 million daltons (10 megadaltons). Adjacent disaccharide units of HAare linked by a β1,4 glycosidic bond, also seen in FIG. 3. Each glcAresidue has a carboxylic acid group (CO₂H) attached to the number 5carbon atom of the glucose ring. Under biological conditions, HA is anegatively charged, randomly coiled polymer filling a volume more than1,000 times greater than would be expected based on molecular weight andcomposition alone. As noted above, the strong negative charges attractcations and water, which allow HA to assume the form of a stronglyhydrated gel in vivo, giving it a unique viscoelastic andshock-absorbing property. HA represents a readily available anddesirable scaffolding material for tissue engineering applications as itis non-immunogenic, non-toxic and non-inflammatory. Also as a naturallyoccurring extracellular matrix (ECM) molecule it offers the advantagesof being recognized by cell receptors, of interacting with other ECMmolecules, and of being metabolized by normal physiological pathways.

Tyramine is a phenolic molecule having an ethyl amine group attachedpara to the OH group on the benzene ring. When these species are used,the mechanism for tyramine substitution onto the singly bound oxygenatom of a CO₂H group on HA proceeds via the carbodiimide-mediatedreaction mechanism described above as illustrated immediately below. Thepreferred carbodiimide species is1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC) as shown.

-   -   where:    -   Structure A is EDC;    -   Structure B is hyaluronan (though only one CO₂H group is shown);    -   Structure C is the product of Reaction A and is        1-ethyl-3-(3-dimethylaminopropyl) isourea;    -   Structure D is tyramine;    -   Structure E is tyramine-substituted hyaluronan; and    -   Structure F is 1-ethyl-3-(3-dimethylaminopropyl)urea (EDU).

In the above pathway, a negatively charged oxygen atom of the carboxylgroup of the hyaluronan molecule attacks, via a nucleophilic reactionmechanism, the electron-deficient diamide carbon atom on thecarbodiimide molecule (EDC) to form the activated O-acylisourea(Reaction A). The result is that the carbon atom of the HA carboxylategroup becomes sufficiently electron deficient to be susceptible tonucleophilic attack by the unshared pair of electrons on the amine groupof a tyramine molecule (Reaction B). Reaction A is preferably catalyzedby a suitable catalyst that will result in the formation of an activeester during Reaction A, thus permitting the reaction to be carried outat substantially neutral pH (e.g. pH=6.5). Suitable catalysts includeN-hydroxysuccinimide (NHS), less preferably 1-hydroxybenzotriazole(HOBt) or N-hydroxysulfosuccinimide (NHSS), less preferably anothersuitable catalyst or combinations thereof effective to enhance thecarbodiimide reaction by formation of an active ester in order tominimize the unproductive hydrolysis of carbodiimides at higher pHs.Less preferably other carbodiimides besides EDC can be used, including1-cyclohexyl-3-[2-(4-methylmorpholino)ethyl]carbodiimide (CMC), anddicyclohexylcarbodiimide (DCC).

The result of Reaction A above is O-acylisourea-substituted hyaluronan;essentially the EDC molecule has been temporarily substituted onto thecarboxylic acid group of a glcA residue from the HA molecule, making thecarbon atom of the carboxylic acid group slightly positively charged.The electron pair from the terminal amine group of a tyramine moleculeis then substituted onto the carbon atom via a nucleophilic substitutionreaction as explained in the preceding paragraph (Reaction B). Theresult of Reaction B is the tyramine-substituted HA molecule (T-HA) andacylurea, a byproduct. It will be understood that Reactions A and B willresult in a plurality of tyramine substitutions on the periodic glcAresidues of HA molecules; a single substitution has been shown here forbrevity and clarity.

After formation of T-HA, a plurality of T-HA molecules are reacted viaperoxide and peroxidase enzyme to cross-link T-HA molecules aspreviously described and illustrated above. That is, the hydroxyphenylgroups on the tyramine residues now attached to HA molecules react withperoxide (preferably H₂O₂) in the presence of a peroxidase to remove thephenolic hydrogen atom resulting in a tyramine free radical, with theunpaired electron associated with the phenolic oxygen atom. This freeradical species isomerizes or resonates, resulting in a resonancestructure (or free radical isomer) with the unpaired electron nowassociated with an ortho carbon atom on the phenolic ring. In thisposition, the unpaired electron quickly reacts with a similarly situatedunpaired electron on another tyramine free radical to form a covalentbond therebetween. The result is a free-radical driven dimerizationreaction between different tyramine free radical residues attached todifferent glcAs of the same or different HA molecules. This dimerizedspecies further enolizes to restore the now-linked tyramine residues,resulting in a dityramine linkage structure. It will be understood thata plurality of reactions as herein described will occur between adjacenttyramine residues, resulting in a cross-linked macromolecular network ofT-HA molecules having the following cross-linking structure:

The cross-linked T-HA network can be provided with aggrecan molecules ina conventional manner, e.g. via link proteins, to provide a cross-linkedT-HA network having aggrecan molecules attached to the HA chains. Thus,a network similar to that found in a normal cartilage aggregate can beprovided, with the dityramine bonds holding the network together therebyconstraining the contained aggrecan network, instead of collagen fibrilsas in normal cartilage.

It will be understood from the present invention that otherglycosaminoglycans (GAGs), polysaccharides and polycarboxylic acids canbe used as the macromolecules for producing the cross-linked networkdisclosed herein. For example, suitable GAGs, other than HA, includechondroitin, chondroitin sulfate, dermatan sulfate, heparan sulfate andheparin. Other suitable polycarboxylates include: proteoglycans such asversican, aggrecan, and cartilage aggregates composed of aggrecan,hyaluronan and link protein; polyuronic acids such as polypectate(polygalacturonic acid), polyglucuronic acid, pectin (polygalacturonicacid methyl ester), colominic acid (poly[2,8-(N-acetylneuraminicacid)]), and alginate (poly[mannuronate-co-guluronate]); and amino acids(having at least 2 amino acid units) that meet the definition ofpolycarboxylate given above, such as polyaspartic acid, and polyglutamicacid. All of these can be substituted with one or a plurality ofhydroxyphenyl groups using the carbodiimide-mediated reaction pathwaydisclosed herein by a person of ordinary skill in the art without undueexperimentation.

As mentioned above, it is also to be understood that native polyphenolcompounds, which already contain two or more hydroxyphenyl groups thatcan be cross-linked using the described enzyme catalysis chemistry canbe used in place of the polycarboxylates and polyamines described abovewhich must have the hydroxyphenyl groups added by a chemical reaction.

In another preferred embodiment, a network of tyramine cross-linkedchondroitin sulfate molecules (either alone or provided as part ofaggrecans) is provided to simulate or replace normal cartilage.Chondroitin sulfate is identical to hyaluronan except: 1) the repeatdisaccharide structure contains N-acetylgalactosamine (galNAc) ratherthan glcNAc, a difference in only the position of the hydroxyl groupattached to the 4-carbon (circled in FIG. 3); 2) the presence ofO-sulfation on the hydroxyl groups at the 4- and/or 6-position of thegalNAc residue and/or the 2-position of the glcA residue (FIG. 3); and3) the size of the chondroitin sulfate chains, which are smaller thanhyaluronan with between 20 to 100 repeating disaccharide units. (Anaggrecan molecule is made up of multiple—roughly 100—chondroitin sulfatechains linked to a core protein through a linkage saccharide located ateach chain's reducing end). In this embodiment, the negatively chargedSO₄ ²⁻ groups of adjacent (cross-linked) chondroitin sulfate moleculesprovide the principal repulsive force contributing to the compressionresistance of the network aggregate while the tyramine cross-linksconstrain the chondroitin sulfate network from breaking or dissipating.The result is a similarly non-displaceable chondroitin sulfate network(and concomitant water-impermeability) as in normal cartilage, butwithout the extracellular collagen fibril matrix or the HA chains foundin normal cartilage. In fact, by directly cross-linking chondroitinsulfate molecules, (instead of their core HA molecules as in thepreviously described embodiment), the repulsive force between adjacentchondroitin sulfate molecules may be strengthened, resulting in evenstronger fluid flow resistance compared to normal cartilage. This mayresult in greater loading force absorption and dissipation capacity thannormal cartilage because the interstitial fluid phase is even moreconstrained from flowing. In this embodiment, where chondroitin sulfatemolecules are directly cross-linked, certain cartilage degenerativeconditions are entirely circumvented; e.g. conditions where the coreprotein to which chondroitin sulfate molecules are ordinarily bonded innormal cartilage becomes cleaved between the HA binding domain (G1) andthe second globular domain (G2) thus allowing the chondroitin sulfaterich region to diffuse out from the cartilage aggregate. In thisembodiment, because the chondroitin sulfate molecules are directlycross-linked to one another, unassociated with an aggrecan or otherproteoglycan molecule, they cannot be cleaved or carried away as innormal cartilage.

Nonetheless, a tyramine cross-linked T-HA network (having an HA backbonechain with attached aggrecan molecules, which in turn includechondroitin sulfate chains) may be preferred because of the highavailability of HA. This may be beneficial in the case of cartilagereplacement or repair using the present invention, because the body'snormal metabolic pathway for generating cartilage may be able to builddirectly onto an implanted tyramine cross-linked T-HA network as will bedescribed.

One further particular application where a cross-linked networkaccording to the invention will have substantial utility is in theproduction of an artificial kidney. The kidney filters blood by twomechanisms: one is by size exclusion and the second is by chargeexclusion. MEMS devices have been designed for use in artificial kidneydevices, which contain precisely defined micropores that can effectivelymimic only the size exclusion characteristics of the kidney. In ahealthy kidney, the charge exclusion related filtration is the result ofheparan sulfate proteoglycans present in a basement membrane, whichseparates two distinct cell types important for other kidney relatedfunctions. To mimic this charge barrier in the MEMS engineeredartificial kidney, hydrogels can be prepared composed of either heparansulfate or heparin that are cross-linked via dihydroxyphenyl(dityramine) links as described herein and provided within the pores ofthe MEMS device. This heparin/heparan sulfate hydrogel can then besandwiched between two hyaluronan derived hydrogels (e.g. T-HA describedabove) as described herein, and containing one of each of the cell typesnormally found in a normally functioning kidney. The centralheparin/heparan sulfate hydrogel provides the charge exclusionproperties for the device. The outer two hyaluronan hydrogel layersprovide protection from the immune system and fouling by normal cellularand molecular debris. Inclusion of the two cell types on opposite sidesof the filtration barrier provides a cellular component in its normalphysiologic orientation.

In another promising application, the hydrogels herein described can beapplied in developing an artificial pancreas. A problem in developmentof an artificial pancreas is the short half life of MEMS engineeredglucose sensors due to fouling of the detector electrode in vivo.Coating of the surface of these detectors with a hyaluronan hydrogel(e.g. T-HA) as described herein would permit diffusion of the smallmolecular weight glucose molecules that they are designed to detectwhile providing protection from the immune system and fouling by normalcellular and molecular debris.

In summary, it will be evident from the foregoing that macromoleculesuseful as scaffold materials for formation of hydrogels include but arenot limited to polycarboxylates (containing free carboxylate groups),polyamines (containing free primary amine groups), polyphenols(containing free hydroxyphenyl groups) and their copolymers, examples ofwhich have been described above. When polyphenols are used, the firststep in preparing the network described above can be omitted becausepolyphenols already contain multiple or periodic hydroxyphenyl groups.Otherwise, both polycarboxylates and polyamines must have hydroxyphenylgroups added or substituted along their length, preferably via theabove-described carbodiimide reaction pathway. The second step inpreparing the network is to carry out an enzyme driven dimerizationreaction between two hydroxyphenyl groups attached to adjacentmacromolecules (whether polycarboxylates, polyamines or polyphenols) inorder to provide a cross-linked structure. This step is carried outusing a peroxide reagent (preferably hydrogen peroxide) in the presenceof a suitable enzyme (preferably HRP) under metabolic conditions oftemperature and pH.

In the case of the preferred dityramine cross-linked T-HA network, inthe first step the carboxyl groups on high molecular weight hyaluronan(HA) are substituted with tyramine which introduces reactivehydroxyphenyl groups into the HA molecule. This tyramine substitutionreaction preferably is mediated by the carbodiimide,1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC) with the degree oftyramine substitution on HA controlled by the molar ratios and absoluteconcentrations of tyramine, EDC and HA used in the reaction mix. Excessreagents such as unused tyramine and EDC are subsequently removed bydialysis, allowing isolation and recovery of high molecular weighttyramine-substituted HA (T-HA). The percent tyramine substitution withineach T-HA preparation is easily calculated by measuring: 1) theconcentration of tyramine present in the preparation, which isquantitated spectrophotometrically based on the unique UV-absorbanceproperties of tyramine at 275 nm (see Example 2 below); and 2) theconcentration of total carboxyl groups in the HA preparation, which isquantitated spectrophotometrically by a standard hexuronic acid assay.By this technique, T-HA preparations which contain a percent tyraminesubstitution of only 4-6% have been routinely synthesizedexperimentally. At this level of tyramine substitution, the vastmajority (preferably at least 60, 70, 80, 90, or 95, percent) of the HAmolecule remains chemically unaltered, and therefore biologicallyfunctional. From this formulation of T-HA (i.e. 4-6% tyraminesubstitution) a wide range of biomaterials with a wide range of physicalproperties can be produced by simply varying the concentration of theT-HA used in the second step of the process.

In the cross-linking reaction, solutions of T-HA are cross-linked toform hydrogels through an enzyme (peroxidase) driven reaction, whichcatalyzes the formation of a covalent bond between two tyramine adductson adjacent HA molecules, producing a single dityramine cross-link. Theformation of multiple, e.g. hundreds, of these dityramine cross-linksper HA molecule result in formation of a stable 3-dimensional scaffoldor hydrogel. Addition of very dilute peroxide (preferably H₂O₂) isrequired to initiate the cross-linking reaction as it is the peroxide,not —HA, that is the actual substrate for the peroxidase enzyme. Theproducts of the reaction of the peroxidase enzyme on peroxide are freeradicals that are preferentially taken up by the hydroxyphenyl rings oftyramine resulting in the formation of the dityramine cross-links. Thedityramine linked structures are fluorescent blue (see Example 2), aproperty which is used to both image the hydrogels and to quantify thedegree of cross-linking within the hydrogels. Since the cross-linkingreaction is enzyme driven, the hydrogels can be formed under physiologicconditions, and therefore can be formed in the presence of includedcells or bioactive agents, or directly adjacent to living tissue whilemaintaining cell and tissue viability.

The resulting hydrogels are optically clear with a wide range ofphysical properties depending on the initial T-HA concentration. Forexample, hydrogels formed from T-HA solutions of 6.25, 12.5, 25, 50 and100 mg/ml T-HA have been shown experimentally to have physicalproperties (rigidity, rheology and texture) of a jelly, a gelatin, adough, a resilient rubber-like composition (similar to a rubber ball),and a cartilage-like material respectively—see Example 3. Thesematerials have potential applications in a wide range of clinicalsettings including tissue engineering of both orthopedic (i.e.cartilage, bone, tendon, meniscus, intervertebral disc, etc.) andnon-orthopaedic (kidney, liver, pancreas, etc.) tissues, gene and drugdelivery, coating of non-biological devices for in vivo implantation(i.e. glucose sensors, artificial hearts, etc.), wound repair, biosensordesign, and vocal cord reconstruction.

Advantageous properties of the hydrogels described herein include theability to: 1) provide easy characterization and quality control; 2)integrate with existing tissue matrices; 3) directly incorporate intonewly formed matrices; 4) directly include cells and bioactive factors;5) maintain biocompatibility; 6) control bioresorption; 7) cast easilyinto complicated anatomical shapes (see Example 4 below); and 8) exhibitthe mechanical properties of native tissues such as articular cartilage.

Current biologically-based surgical procedures for cartilage repairinclude autologous chondrocyte implantation, drilling, abrasionchondroplasty, microfracture, and mosaic arthroplasty. All theseprocedures treat only focal articular cartilage injuries, and notcartilage denuded joint surfaces such as seen in severe osteoarthritisand rheumatoid arthritis. Also, they use either cartilage tissue plugsor expanded chondrocytes harvested from the patient to fill cartilagedefects. These tissues or chondrocytes are expected to fill the defectby synthesizing entirely de novo material, such as newly synthesizedhyaline cartilage, that has integrated with existing cartilage matricesand has the biomechanical properties of normal cartilage. However, suchprocedures all promote the formation of a reparative tissue(fibrocartilage) rather than true hyaline cartilage with furthermechanical damage to fibrocartilage thought to predispose the joint toosteoarthritis. Furthermore, the availability of endogenous cartilage asa repair material is quite limited with its acquisition presenting itsown risks and morbidity to the patient. As evident from the foregoingdiscussion and as will become further apparent based on the followingExamples, the synthetic macromolecular networks and resulting hydrogelsdisclosed herein present practical materials for promising new therapiesin patients suffering from cartilage degenerative diseases. Thematerials are entirely synthesized from commercially available ex vivoreagents and so involve no morbidity to the patient which conventionallywould be required to harvest endogenous material. In addition, thehydrogel (particularly T-HA) can be implanted as an effective cartilagesubstitute in cartilage denuded joints as a direct intervention forpatients suffering from cartilage-degenerative diseases because they canbe synthesized so as to emulate the behavior of normal, healthycartilage.

Rather than relying on synthetic or natural materials or on chondrocytesto produce de novo an implantable, synthetic cartilage-likeextracellular matrix (ECM), the present inventors initially focused onpurifying the molecules that give cartilage its form and structuralcharacteristics, and then minimally modifying these molecules to make amaterial resistant to biological degradation. While chondrocytes stillmay be relied on for maintenance of the synthetic ECM provided by themacromolecular (e.g. T-HA) network post-implantation (e.g. chondrocytescan be embedded into the hydrogel materials as described above), theyare not relied on for de novo synthesis. Instead, the basic structure ofthe synthetic materials described here is modified by cross-linking viaa dihydroxyphenyl, preferably dityramine linkage chemistry, to ensureits survival. On further development and experimentation, as will beseen in the following Examples it was discovered that hydrogels can bemade from such materials having a wide array of viscoelastic and otherphysical properties that can be tuned by appropriate and judiciousselection of reagent concentrations and cross-linking conditions toapproximate or simulate the properties of other native tissues for whichit is or may be desirable to provide a synthetic implantable substitute.

As the Examples below demonstrate, the present hydrogels can be preparedhaving widely varying properties that are suitable for any number ofsynthetic tissue implantation or augmentation, as well as other clinicalapplications. As already described, the present materials can be used torepair cartilage defects produced as a result of either injury ordisease. Defects due to injury that can be so repaired can be sports- oraccident-related, and may involve only the superficial cartilage layer,or may include the underlying subchondral bone. Defects due to diseasewhich can be repaired using the compositions described herein includethose resulting from osteoarthritis and rheumatoid arthritis. Whetherfrom injury or disease, such defects may be in either mature or growthplate cartilage. Formulations for hydrogels for synthetic growth platecartilage may require the inclusion of unsubstituted scaffold materialin order to allow for controlled bioresorption of the biomaterial duringgrowth.

Another potential clinical application for treatment of damaged orarthritic joints is as a replacement for synovial fluid. Conventionallyreferred to as viscosupplementation therapy, this currently involvesinjection of a solution of uncross-linked HA into a damaged or arthriticjoint, which provides sustained pain relief for weeks even though the HAis cleared from the joint in 1-2 days. Use of the T-HA hydrogelsdescribed herein should provide an extended benefit due to their longerin vivo persistence compared to uncross-linked HA.

Another field where the hydrogels described herein will be useful is therepair, reconstruction or augmentation of cartilaginous as well as softtissues of the head and neck. The availability of biomaterials for softtissue augmentation and head and neck reconstruction has remained afundamental challenge in the field of plastic and reconstructivesurgery. Significant research and investment has been undertaken for thedevelopment of a material with appropriate biological compatibility andlife span. The outcomes of this research have not been promising. Whenplaced in immunocompetent animals the structural integrity of currentlyproposed materials has been shown to fail as the framework is absorbed.Furthermore, though conventional synthetic materials offer excellentlifespan, they present certain unavoidable pitfalls. For example,silicones have been fraught with concerns of safety and long-term immunerelated effects. Synthetic polymers PTFE (gortex) and silastic offerless tissue reactivity but do not offer tissue integration and canrepresent long term risks of foreign body infections and extrusion. Thematerials described in this application will be useful to prepare asynthetic soft-tissue scaffold material for the augmentation or repairof soft-tissue defects of the head and neck. In particular, across-linked tyramine-substituted hyaluronan (T-HA) hydrogel, which isnon-inflammatory, non-immunogenic, and which can be prepared having theappropriate degree of viscoelasticity (see Examples below), could beused as an effective implantable scaffold material. In addition, theunique ability of the preferred enzyme-driven cross-linking chemistry tomaintain cell viability permits inclusion of cells such as chondrocytesdirectly into the hydrogels during formation which can be performed insitu at a defect site. Thus, the need to sculpt or mold an anatomicallycompatible graft shape to fit a particular defect site is eliminated.

The dityramine cross-linked T-HA network described above has particularutility for producing artificial or synthetic cartilage. The presenthydrogel materials can be used, for example, as a novel, biocompatibleand biocompliant material to prepare cartilage implants which arefrequently used in reconstructive procedures of the head and neck torepair cartilaginous or bony defects secondary to trauma or congenitalabnormalities. Applications specific to the ear include otoplasty andauricular reconstruction, which are often undertaken to repaircartilaginous defects due to trauma, neoplasm (i.e., squamous cellcarcinoma, basal cell carcinoma, and melanoma), and congenital defectssuch as microtia. Applications specific to the nose include cosmetic andreconstructive procedures of the nose and nasal septum. Dorsal humpaugmentation, tip, shield and spreader grafts are frequently used incosmetic rhinoplasty. Nasal reconstruction following trauma, neoplasm,autoimmune diseases such as Wegeners granulomatosis, or congenitaldefects require cartilage for repair. Septal perforations are difficultto manage and often fail treatment. Cartilage grafts would be ideal forthese applications, as autologous or donor cartilage is oftenunavailable. Applications specific to the throat include laryngotrachealreconstruction, which in children usually requires harvesting costalcartilage, which is not without morbidity. Auricular and septalcartilage is often inadequate for this application. Syntheticcartilaginous materials prepared from hydrogels disclosed herein can besynthesized to suit each of the foregoing applications, based on tuningparameters of hydrogel synthesis such as reagent concentration,substitution and cross-linking rates, etc., as evident from the belowExamples. Laryngotracheal reconstruction is usually performed for airwaynarrowing due to subglottic or tracheal stenosis. The etiology may betraumatic (i.e., intubation trauma, or tracheotomy) or idiopathic. Otherpossibilities include chin and cheek augmentation, and use in ectropionrepair of the lower eyelid, in addition to numerous craniofacialapplications. It should be noted that these applications may not needcartilage with the exacting mechanical properties of articularcartilage. Inclusion of a cell population or bioactive agents may alsobe desirable.

The hydrogel materials described herein also can be used for repair andnarrowing of the nasal cavity, normally following overly aggressivesurgical resection, to prevent the chronic pooling of fluid in the nasalpassages that leads to infection and encrustation. Another promisingapplication is in laryngotracheal reconstruction in both children andadults, as a result of laryngotracheal injury due for example tointubation during a surgical procedure such as cardiovascular surgery.Damaged tracheal cartilage at the anterior and posterior portion of thetracheal ring can be replaced with pre-cast hydrogel formed in the shapeof an elongated blocked “T” or an inside out canoe, e.g. via methodsdisclosed below in Example 4. Hydrogels as herein described also can beused:

-   -   to provide cricoid ring replacements.    -   to protect the carotid artery following neck resection for        cancer—the hydrogel can be placed between the carotid artery and        the skin as a protective barrier for the carotid artery against        loss of the skin barrier.    -   as a protective coating during neuronal repopulation of a        resected nerve—often fibrous tissue forms faster than the        neuronal repopulation preventing its eventual formation.        Placement of the nerve ends within a hydrogel pre-cast tube        could exclude fibrous tissue formation from the site of        repopulation.    -   for reconstruction of the mastoid cavity following ablative ear        resection normally as a result of ear infection.    -   for inner ear reconstruction; specifically in place of        prosthetic silastic implants for anvil/stapes replacements. The        hydrogels can be used to replace natural cartilage used as the        top portion of these graphs, or to completely replace these        graphs with an entirely hydrogel graph construct.    -   for repair of soft tissue defects including chin and cheek        augmentation, and use in ectropion repair of the lower eyelid,        in addition to numerous craniofacial applications.    -   for cosmetic and reconstructive purposes in sites other than the        head and neck, for example use as breast implants for breast        augmentation.    -   as a wound sealant, for example to fill the void left after        removal of lymph nodes (i.e. due to cancer) in the breast or        neck, to seal the lymphatics and abate uncontrolled fluid        drainage into the resection site that may lead to infection and        other complications.

In addition to synthetic cartilaginous tissues as described above, themacromolecular network materials described herein and the hydrogels madefrom them also can be used in other tissue engineering applications toproduce other synthetic orthopaedic tissues, including, but not limitedto, bone, tendon, ligament, meniscus and intervertebral disc, usingsimilar strategies and methodologies as described above for thesynthesis of artificial forms of cartilage. As evidenced in the Examplesbelow, the materials also can be used to make synthetic non-orthopaedictissues including but not limited to vocal cord, vitreous, heart valves,liver, pancreas and kidney, using similar strategies and methodologiesas described above for the synthesis of artificial forms of cartilage.

Another field where the hydrogel materials disclosed herein presentpromising utility is in certain gastrointestinal applications where itis necessary to treat or prevent the formation of scar tissue orstrictures in abdominal or gastrointestinal organs. There already are anumber of products at various stages of clinical and FDA approval, whichgenerally are termed ‘hydrogels,’ that are designed or intended to beuseful in the treatment and prevention of scarring and/or strictureformation. The hydrogels of the present invention are superior to theseother hydrogels in that the ones disclosed here can be made entirelyfrom non-immunogenic materials as opposed to exogenous materials such assilicones or other synthetic polymers, and they can be cross-linked insitu within a patient. The hydrogel compositions disclosed herein can beused in similar applications as the already known hydrogels are used orintended to be used, including the following:

-   -   for treatment of strictures or scarring of the gastrointestinal        tract. The treatment involves injection of the hydrogel material        at the site of an anticipated stricture to prevent scarring, or        at a site of existing stricture after therapy to enlarge the        narrowed GI tract to prevent the stricture from reoccurring.    -   for treatment of esophageal strictures. Esophageal strictures        are a common complication of gastroesophageal reflux disease        (GERD). GERD is caused by acid, bile and other injurious gastric        contents refluxing into the esophagus and injuring the        esophageal lining cells. Approximately 7-23% of GERD patients        develop an esophageal stricture, or fibrous scarring of the        esophagus. Esophageal scarring also can be caused by ablative        therapies used to treat Barrett's esophagus. The major        complication of such ablative therapies is that the ablative        injury extends too deeply into the esophageal wall and results        in an esophageal scar or stricture. Esophageal strictures        prevent normal swallowing and are a major cause of patient        morbidity. The hydrogel materials described herein may be used        to treat or prevent esophageal strictures resulting from GERD,        Barrett's esophagus, and esophageal ablative therapies.    -   for treatment of Crohn's disease. Crohn's disease causes        strictures or scars that block off or narrow the lumen of the        bowel, preventing normal bowel function. The present hydrogels        may be useful to treat or prevent such strictures.    -   for treatment of primary sclerosing cholangitis (PSC). PSC is a        rare disease of the bile ducts of the liver. The bile ducts form        a branching network within the liver and exit the liver via two        main branches that are combined into the common bile duct which        drains the liver and gallbladder of bile into the duodenum. The        bile ducts are very narrow in diameter, measuring only up to 2        mm normally at their largest most distal portions, and yet they        must normally drain liters of bile every day from the liver into        the duodenum. Any blockage of these ducts can result in a        serious condition known as jaundice, which allows many toxins        and especially hemoglobin breakdown products to accumulate in        the body. PSC is a scarring or structuring disease of the bile        ducts within the liver and in the extrahepatic bile ducts        described above that connect the liver to the small intestine.        The bile duct strictures of PSC may be treated or prevented with        the present hydrogels.    -   for treatment of chronic pancreatitis. Chronic pancreatitis is a        chronic inflammatory disease of the pancreas that may be        complicated by scars or strictures of the pancreatic ducts.        These strictures block the drainage of pancreatic juice, which        normally must exit the pancreas through a system of ducts or        drainage conduits into the small intestine. The pancreatic juice        contains many digestive enzymes and other elements important to        normal digestion and nutrient absorption. Blockage or narrowing        of the pancreatic ducts by chronic pancreatitis can results in        severe complications in which the pancreas autodigests and forms        life-threatening abdominal infections and or abscesses. The        pancreatic strictures of chronic pancreatitis may be treated or        prevented with the present hydrogels.    -   for treatment of gallstone-induced bile duct and pancreatic duct        strictures. Gallstones are a very common disorder, a principal        complication of which is the formation of bile duct and        pancreatic duct strictures, which may be treated or prevented        with the hydrogels.    -   for treatment of ischemic bowel disease. The intestines are        prone to the formation of scars or strictures when their blood        supply is compromised. Compromised blood flow is called        ischemia, and can be caused by many pathologies, including        cardiovascular disease, atherosclerosis, hypotension,        hypovolemia, renal or hepatic disease-induced hypoalbuminemia,        vasculitis, drug-induced disease, and many others. The end stage        result of all of these etiologies can result in intestinal        strictures that block off the bowel and prevent its normal        function. The present hydrogels may be used to treat or prevent        ischemic bowel strictures.    -   for treatment of radiation-induced intestinal strictures.        Radiation therapy for cancer is associated with numerous        morbidities, important among which is intestinal stricture        formation. The present hydrogels may be used to treat or prevent        radiation-induced intestinal strictures.

In addition to making synthetic tissues, the hydrogels disclosed herealso can be used to provide a coating for non-biological structures ordevices to be used in surgery or otherwise for in vivo implantation,such as surgical instruments, or ceramic or metal prostheses. Such acoating would provide a barrier between the non-biologic device materialand living tissue. The role of hydrogels as a barrier for non-biologicdevices includes, but is not limited to: 1) prevention of absorption ofmacromolecules and/or cells on the surfaces of non-biologic devices,which can lead to protein fouling or thrombosis at the device surface;2) presentation of a non-toxic, non-inflammatory, non-immunogenic,biologically compatible surface for devices made from otherwisenon-biologically compatible materials; 3) compatibility with devicefunction such as diffusion of glucose for a glucose sensor, transmissionof mechanical force for a pressure sensor, or endothelization of avascular graft or stent; 4) enhancement of device function, such asproviding a charge barrier to an existing size barrier in a MEMS basedartificial nephron; 5) incorporation into non-biologic devices of aviable cell population entrapped within an aqueous, physiologicallycompatible environment; and 6) inclusion of drugs or bioactive factorssuch as growth factors, anti-viral agents, antibiotics, or adhesionmolecules designed to encourage vascularization, epithelization orendothelization of the device.

Based on the foregoing, the hydrogels of the present invention may beused to provide a non-allergenic coating for a variety of implantabledevices including an implantable glucose sensor for management ofdiabetes. In addition, the hydrogels may be used to provide: a chargebarrier for the development of MEMS-based artificial nephrons; anaqueous, physiologically compatible environment in which embedded kidneycells such as podocytes can be incorporated into a MEMS-based artificialnephron design; and a coating for implantable MEMS devices designed fora variety of purposes including, but not limited to, drug delivery,mechanical sensing, and as a bio-detection system.

The disclosed hydrogels, and particularly a hyaluronan-based hydrogel,also may be covalently attached to silicon-based devices, e.g. throughfirst covalent attachment of the primary amine of tyramine to thesilicon surface to provide a hydroxyphenyl coated surface chemistry.This may use the same chemistry used to bind DNA that has been modifiedwith a free amine to silicon surfaces. The HA-based hydrogel then iscovalently coupled to the hydroxyphenyl coated surface by the sameperoxidase driven chemistry used in its preferred cross-linking modedescribed above.

The hydrogels also can be used for coating non-biologic cardiovasculardevices such as catheters, stents and vascular grafts. These wouldinclude devices made from materials conventionally not used because oftheir biological incompatibility, but which have superior designcharacteristics to those devices currently in use. Bioactive factorscould be incorporated into the hydrogels to promote endothelization orepithelization of the hydrogel, and thus of the implanted device.

A particularly promising application mentioned above is in the designand implementation of an implantable artificial glucose sensor for thetreatment and management of diabetes. Effective glycemic controlrequires frequent measurement of blood glucose levels, which currentlyrequires a pin prick (or “finger stick”) to obtain a blood sample. Thereis tremendous clinical interest in a reliable, cost-effective method ofblood glucose measurement and in preventing hypoglycemia, which is thecause of most severe life-threatening events. From a technologicalstandpoint, microsensors have been very successful over the last decadein a wide variety of applications. The successful development of abiocompatible long term implantable glucose sensor would significantlyimpact routine monitoring of glucose levels by diabetic individuals andplay a major contributory role in the further development of abioartificial pancreas.

A design of a sensor for use during cardiovascular surgery has beenpublished, Clark L C, Lyons C, “Electrode system for continuousmonitoring in cardiovascular surgery,” Annals of New York Academy ofScience, 102:29-45 (1962). Subsequently, efforts have been directedtoward developing and testing an implantable device that could mimic thenative glucose/insulin control system. Besides the obvious advantage ofserving as part of a bioartificial pancreas, such a system could becoupled with telemetry hardware and thereby give the patient advancewarning of hypoglycemia.

Prior work on implantable glucose sensors generally follows one of twoapproaches. The first involves placing sensors into blood vessels suchas the vena cava or the carotid artery. The second involves placingsensors subcutaneously. These sensors may involve a microdialysis probeor more commonly, an amperometric enzymatic-based transducer. It isbelieved the risk of thrombosis and hematogenous spread of infectionmitigate against the long term use of intravascular sensors. While theexact relationship between blood and subcutaneous glucose concentrationsis still being investigated, recent work suggests that mass transfermodeling methods can significantly improve the estimates of bloodglucose levels that are based on subcutaneous data. Furthermore, thereare significant advantages associated with subcutaneous sensors:clinical safety, ease of insertion and removal, ease of coupling thesesensors to a telemetry system and cost. There is substantial evidencethat subcutaneous placement of a glucose sensor will work and will leadto much longer life of the sensor than if it were to contact blooddirectly.

However, a major problem in the design of any continuous glucose sensorfor clinical use remains the long-term drift of the sensor, usuallycaused by fouling of the electrode when exposed to human tissue or thegradual loss of enzyme activity. The introduction of various membranesto act as a glucose or a hydrogen peroxide barrier has, in general,improved sensor performance but it has not resulted in long termstability. The much heralded membrane for this purpose, Nafion, rapidlydeteriorates when implanted in the body. Introduction of an implant intosubcutaneous tissue elicits both acute and chronic inflammatoryresponses. Together these result in a complexly orchestrated growth ofnew tissue which ultimately envelops the implant with a foreign bodycapsule (FBC). In the short term, it is likely that inflammatory cellsmetabolize glucose and thereby cause artifacts in the glucose readings.When discussing the problems with long-term use of subcutaneous sensors,experts maintain that the diminished response in vivo can be ascribed tothe protein or cellular coating around the sensor which interferes withthe mass transport of glucose. If suitable covering membranes for thesensor could be provided to exclude interfering substances or controlcoating or encapsulation with proteins and cells, the excellentperformance in vitro may be matched in vivo. The use of the HA-basedhydrogels described herein as a coating agent to both minimize the FBCand keep it away from the sensor membrane should prove a usefulsolution.

The purpose is to control the tissue response to an implantable glucosesensor using a HA-coating on the sensor membrane. A sheath of HA-basedhydrogel will give the sensor membrane “breathing space” by preventingproteins and cells from clogging the diffusion of glucose and oxygeninto the sensor. Prior experience has indicated that HA and itsderivatives are extremely biocompatible and as a consequence are used insituations where the host tissue response needs to be minimized (e.g.,in eye implantation surgery). Thus, sensor performance should beenhanced in the long term when HA-based hydrogels are cast around sensormembranes as it relates to the development of an implantable glucosesensor with the long term perspective that such a sensor should resultin improved blood glucose monitoring and ultimately improved quality oflife for the diabetic population. In addition, the novel cross-linkingstructure of the HA-based hydrogels herein disclosed will ensurelong-term maintenance of such a coating which will provide significantlongevity to a subcutaneously implanted glucose sensor.

Still another promising application is in the production of abioartificial kidney for the treatment of end-stage renal disease(ESRD). The only current treatment options for ESRD patients are renalreplacement therapy (all forms of dialysis) and transplantation.Transplantation is limited by the shortage of donor organs, and iscomplicated by the necessary and expensive life-long use ofimmunosuppressive drugs. Alternatively, although dialysis can prolongthe life of ESRD patients, average life expectancy on dialysis isreduced by 50%, and the remaining quality of life is far from ideal.Repeated vascular access and handling of the patient's blood leads tofrequent and sometimes life threatening infections.

The functional unit of the kidney is the nephron. The nephron beginswith a filtering structure, the glomerulus, which is of a tuft ofcapillaries surrounded by epithelial cells (podocytes) and supported bymesenchymal cells (the mesangium). The glomerulus is connected directlyto the tubule of the nephron, a long tube lined with a single layerepithelium of polarized cells. The tubule cells function to salvagefluid, electrolytes and nutrients from the filtrate (by bothintracellular transport and pericellular movement) concentrating thefiltrate into urine. All nephrons connect into the collecting system, anetwork of epithelial-lined tubes, which has some additionalreabsorptive properties, but primarily functions to direct the urine tothe bladder. The filtration unit of the nephron, the glomerulus,consists of the endothelial cell of the capillary arteriolar wall, thepodocyte surrounding the exterior of the capillary, and the glomerularbasement membrane (GBM) sandwiched between the two cell types. Theglomerular capillaries are some of the smallest vascular beds in thebody, and the glomerular endothelial cells are specialized for theirfunction by being fenestrated to allow direct contact of the bloodplasma to the filtration barrier. Although these fenestrated endotheliado restrict the movement of leukocytes and very large molecules into thefiltrate, the permselectivity of the filtration barrier is defined bythe podocyte and the GBM.

The GBM is a classic basement membrane structure composed of theprototypic molecules: type IV collagen (α3, α4, α5 heterotrimers),laminin (Laminin-11, α5, β2, γ1 heterotrimers), HS proteoglycans(perlecan and agrin) and nidogen (nidogen-1 and -2); as well as severaladditional ECM molecules including, collagen V, fibronectin, a CSproteoglycan (bamacan) and several small leucine-rich proteoglycans(biglycan, decorin, podocan). The GBM is synthesized by both theendothelial cell and the podocyte. Each cell produces a completebasement membrane which subsequently fuses during development to formone, double thickness basement membrane. The GBM has important functionsin providing the appropriate microenvironment and substrata for thepodocytes and endothelial cells. Without a normal GBM, both cell typeslose their typical morphology and cellular differentiationcharacteristics, which subsequently destroys glomerulus function. TheGBM also functions in filtration by restricting the movement of waterand has some contribution to the size and charge selectivity, however,the majority of permselectivity is dictated by the podocyte.

The podocyte is a highly specialized epithelial cell and has uniquefunction in the glomerulus. The podocyte extends lamellipodia that wraparound the capillaries, branching into very fine interdigitations withother podocytes. On cross section, these interdigitated cellularextensions are called foot processes (FP) and the spaces between theFPs, where filtration occurs, are called slits. The podocyte synthesizesa macromolecular structure that spans the slit, the slit diaphragm (SD),which forms a bridge between two adjacent FPs. The molecular compositionand structure of the SD is not fully understood. The SD appears to be amodified adherens junction containing additional podocyte-specificproteins, the most notable being nephrin. Nephrin extends from theplasma membrane of one FP and forms a homodimeric interaction withanother nephrin molecule extending from the adjacent FP, creating azipper-like structure when viewed in cross section by electronmicroscopy. How the SD and nephrin function as a permselective barrieris not known, but is currently a very active area of research.

Biological microelectromechanical systems (bioMEMS) are a promising areaof exploration for development of next generation bioartificial kidneys.Drug delivery systems, immunoisolators, and capillary networks, as wellas precise control of cell differentiation and growth have beendemonstrated for bioMEMS. The kidney is the first organ for whichchronic substitutive therapy has been accepted, and application of thebioMEMS toolkit to treatment of ESRD is both evolutionary in thetechnology and revolutionary in the end product. Silicon micromachiningtechnology has evolved such that structures with feature sizes on theorder of 1-100 nanometers can be reliably produced in quantity. Thesedimensions are on the order of those for the glomerular slit diaphragm.The facility with which standard silicon bulk and surface micromachiningtechnology permits microfluidic control, patterned deposition of cellsand extracellular matrix proteins, and immunoisolation of cells, lendsitself to tissue engineering of artificial organs. The engineering ofnanoscale semiconductor filtration membranes could permit independentcontrol and investigation of charge-size selectivity with the potentialto lead to the tissue engineering of a bioartificial glomerulus andeventually a complete nephronal unit.

One of the first components in the miniaturization of a bioartificialkidney is development of a nanofabricated hemofiltration membrane (NHM)from bioMEMS components. The NHM is intended to serve the hemofiltrationfunction of the glomerulus in the nephron-like devices of abioartificial kidney. NHM arrays can be fabricated using standardsilicon micromachining techniques containing slit pores of approximatelythe dimensions of the glomerular slit diaphragm, and conductedexperiments to demonstrate its size barrier characteristics similar tothose of the glomerular basement membrane. The chemistries and hydrogelsdescribed in this patent application can be used to provide twoadditional and necessary components to the filtration characteristics ofthe NHM that are required for glomerular function. The first is a chargebarrier component similar to those of the glomerular basement membrane.This would be provided by application of a layer of heparan sulfate (HS)based hydrogel. HS is a type of GAG similar to HA and CS. The secondaddition is inclusion of the podocytes, which are responsible for themajority of the filtration function of the glomerulus through the slitdiaphragm. The podocytes would be applied to the surface of the HS-basedhydrogel layer in a HA-based hydrogel layer, which would also serve toprovide a layer of biocompatibility. The presence of the HS layer shouldfacilitate proper matrix-cell interactions and stimulate the depositionof an appropriate basement membrane.

The hydrogels described herein, including but not limited totyramine-based hyaluronan hydrogels, also can be used as research andclinical reagents. One promising application is controlled or extendedrelease drug delivery. In this application, the drug can be trappedwithin a sphere or other suitable shape of hydrogel material composed ofa central spherical or other shaped core of hydrogel formulated at arelatively high macromolecular concentration (and thus lowest porosity),onto which concentric spherical layers of hydrogel are coated, eachsuccessively coated layer being formulated of a progressively lowermacromolecular concentration (and thus higher porosity). Release of thedrug is then controlled by the rate of hydrogel degradation if soengineered, binding of the drug to the hydrogel scaffold and diffusionof the drug through the scaffold pores. The hydrogel sphere then isimplanted into a patient at an appropriate location to effect extendedrelease of the drug.

Targeted drug deliver also can be achieved through an affinity-basedstrategy based on designed affinity of drug laden hydrogel particles tospecific tissue and cell types. To this end, the hydrogels can be usedas an affinity-based medium for the selective binding and thuspurification of specific cell populations through incorporation oftargeted cell binding molecules within the hydrogels during or prior tocross-linking. Once a select cell population is bound to the hydrogelaffinity-based medium they could be released for further investigation,or directly entrapped while bound to the hydrogel affinity-based mediuminto other formulations of the hydrogels for other tissue engineering orclinical applications.

Such an affinity-based medium also can be used for the selective bindingand purification of hyaluronan binding proteins. As the entire mediumcan be made solely of hyaluronan with no other support materialbackground binding should be quite low. By using other materials as thescaffold material (such as aggrecan) other affinity-based media can beprepared for purification of molecules that selectively bind to thosescaffold materials.

Such an affinity-based medium also can be used for selective binding andpurification of specific macromolecules or cell populations throughincorporation of protein A within the hydrogels during cross-linking.Antibodies specific to the macromolecule or cell population of interestcan then be used to coat the protein A infused hydrogels with theantibodies optimally oriented with their antigen binding (F_(ab)) armsdirected outward and their constant (F_(c)) domain bound to the proteinA. Once a select cell population or macromolecule is bound to theprotein A hydrogel, it could be released for further investigation, ordirectly entrapped while bound to the protein A hydrogel into otherformulations of the hydrogel for other tissue engineering or clinicalapplications. Alternatively antibody could be directly incorporated intothe hydrogels.

The disclosed hyaluronan-based hydrogel materials also have utility as adiagnostic for the presence of hyaluronidases which can be predictive ofthe metastatic potential of certain cancers; e.g. by coating of a biopsyslide with hyaluronan hydrogel and measurement of the extent andlocalization of the loss of intrinsic fluorescence of the hydrogelmaterial due to its dityramine cross-links as the hydrogel is digestedby endogenous hyaluronidases. By using other materials as the scaffoldmaterial (such as aggrecan) other degradative enzymes could be detectedsuch as metaloproteinases.

Further aspects of the invention will be understood in conjunction withone or more of the following examples, which are provided by way ofillustration.

EXAMPLES Example 1

Experimental quantities of tyramine-substituted hyaluronan hydrogelshaving dityramine cross-links according to the invention have beenprepared as follows. HA is dissolved at 1 mg/ml based on hexuronic acidin 250 mM 2-(N-morpholino)ethanesulfonic acid (MES), 150 mM NaCl, 75 mMNaOH, pH 6.5 containing a 10-fold molar excess of tyramine relative tothe molar concentration of HA carboxyl groups. Tyramine substitutiononto the carboxyl groups is then initiated by the addition of a 10-foldmolar excess of EDC relative to the molar concentration of the HAcarboxyl groups. A 1/10th molar ratio of N-hydroxysuccinimide (NHS)relative to the molar amount of EDC is added to the reactions to assistthe EDC catalyzed amidation reaction by formation of active esters.Reactions are carried out at room temperature for 24 hours, after whichthe macromolecular fraction is recovered from unreacted small molecularweight reactants such as tyramine, EDC, NHS, and MES by exhaustivedialysis versus 150 mM NaCl and then ultrapure water followed bylyophilization. After lyophilization, the tyramine-substituted HA (T-HA)product is dissolved to working concentrations of between 5 and 100mg/ml in PBS (which is a buffer compatible with cell suspension, in vivotissue contact, and the cross-linking reaction) to provide variousconcentration preparations depending on the desired rigidity of thefinal hydrogel. Alternatively, the solvent can be any other suitablesolvent besides PBS that will not substantially negatively impact theenzyme activity and that will not interfere with cross-linking reactionvia selective uptake of free radicals generated by the enzyme. Suitablealternative solvents include water, conventional biological tissueculture media, and cell freezing solution (generally composed of about90% blood serum and about 10% dimethyl sulfoxide). Prior to suspensionof cells (see Example 5) or contact with tissues in vivo, the T-HAshould be filtered through a 0.2 μm filter. Next, tyramine-tyraminelinking is carried out by adding 10 U/ml of type II horseradishperoxidase (HRP) to each T-HA preparation. Cross-linking is initiated bythe addition of a small volume (1-5 μl) of a dilute hydrogen peroxidesolution (0.012%-0.00012% final concentration) to yield the finalhydrogel with desired rigidity. For preparation of larger quantities orvolumes of a desired hydrogel, quantities of reagents provided in thisparagraph could be scaled up appropriately by a person of ordinary skillin the art.

Example 2

An experiment was conducted to determine the degree of tyraminesubstitution (and consequent dityramine cross-linking) for a T-HAmacromolecular network according to the invention. Initially, threeformulations of (uncrosslinked) tyramine-substituted hyaluronan (T-HA)were prepared as described above, designated 0×, 1× or 10×. The 0×formulation was prepared using no EDC (i.e. containing no carbodiimide),meaning there was no carbodiimide present to mediate the reaction forcreating an amide bond between the NH₂ group on tyramine and a CO₂Hgroup on the HA molecules. Thus, the 0× formulation can be considered acontrol. The 1× formulation contained a 1:1 stoichiometric ratio of EDCbased on the quantity of CO₂H groups present on the HA molecules in thereaction mixture. The 10× formulation contained a 10:1 stoichiometricratio (or 10-fold excess) of EDC based on the quantity of CO₂H groupspresent on the HA molecules in the reaction mixture. In all threeformulations, a stoichiometric excess of tyramine was provided relativeto the quantity of CO₂H groups on HA. In all three formulations (0×, 1×and 10×) the reactants and the appropriate amount of EDC for theformulation were combined in a vial and agitated to facilitate thetyramine-substitution reaction. All three formulations were allowed toreact for 24 hours at room temperature, after which the vial contentswere dialyzed to remove unreacted tyramine molecules, EDC and acylurea(EDU) byproducts of the reaction. These molecules were easily separatedfrom HA and any formed T-HA molecules through dialysis due to therelatively small size of tyramine, EDC and EDU compared tomacromolecular HA. Once unreacted tyramine and EDC were removed, theremaining contents for each formulation were analyzed to determine therate of tyramine substitution relative to the total number of availableCO₂H sites present on HA molecules.

Tyramine exhibits a UV absorbance peak at 275 nm, making the degree oftyramine substitution easily detectable against a tyramine calibrationcurve. Based on UV-spectroscopic analysis of the above three T-HAformulations, it was discovered that the HA-tyramine substitutionreaction carried out with no EDC present (formulation 0×) resulted insubstantially zero tyramine substitution onto the HA molecules. Thisconfirmed the importance of using a carbodiimide reaction pathway in thetyramine substitution reaction. However, the tyramine absorption in theT-HA formulation prepared using a 1:1 EDC:CO₂H stoichiometric ratio inthe tyramine substitution reaction (formulation 1×) resulted in atyramine substitution rate of about 1.7% relative to all available CO₂Hgroups on the HA chains. The 10× formulation (10:1 EDC:CO₂H ratio)resulted in about a 4.7% substitution rate.

Subsequently, hydrogen peroxide and horseradish peroxidase (HRP) wereadded to each of the three dialyzed HA/T-HA formulations (0×, 1× and10×) at 5 mg/mL and the resulting formulations were allowed to react tocompletion. After reaction in the presence of peroxide and HRP, it wasobserved that the 0× formulation remained entirely liquid, having astrong meniscus; no gel formation was observed, confirming the fact thatno or substantially no tyramine substitution had occurred when no EDCwas used in the tyramine substitution reaction. For the 1× formulation,only a very weak meniscus was observed and the contents of the vial hadgelled, confirming that both tyramine substitution and cross-linking hadoccurred. For the 10× formulation, a relatively rigid gel had formed,and in fact had shrunk relative to the initial volume of fluid in thecontainer, leaving a quantity of liquid (having a meniscus) on top. Thegel prepared from the 10× formulation (having a 4.7% tyraminesubstitution rate) was much firmer and more rigid than that from the 1×formulation having a 1.7% tyramine substitution rate.

The dityramine structure exhibits a blue fluorescence on exposure to UVlight. The products of each of the above formulations were exposed to UVlight to detect the presence of dityramine cross-links; As expected,both the 1× and 10× hydrogels exhibited blue fluorescence (the 10×hydrogel fluorescence being more intense than that of the 1× hydrogel),while the 0× formulation exhibited no blue fluorescence at all. Thisconfirmed the presence of dityramine cross-links in both hydrogels, andthat the occurrence of dityramine in the more rigid hydrogel (10×) wasgreater than in the less rigid hydrogel (1×).

The overall result was that the importance of the carbodiimide-mediatedreaction pathway was demonstrated, and it was confirmed that therelative rigidity of a hydrogel formed from a cross-linked T-HA networkis proportional to the degree of dityramine cross-linking, which is inturn proportional to the degree of tyramine-substitution onto HA. It wasquite a surprising and unexpected result that even a 1.7%tyramine-substitution rate (and subsequent cross-linking rate to formdityramine links) provided a suitably firm T-HA gel (or hydrogel). A4.7% substitution (and cross-linking) rate resulted in even a firmerT-HA gel. Also surprising was that a ten-fold stoichiometric excess ofcarbodiimide (EDC) relative to the quantity of carboxylic acid groupspresent in the reaction mixture (formulation 10×) resulted in only abouta 4.5-4.7% tyramine substitution rate, yet stable and cohesive tyraminecross-linked T-HA networks were nonetheless achieved.

This means that the majority of the carboxylic acid groups on the HAmolecules are unsubstituted and not tyramine cross-linked, essentiallyremaining the same as in the native HA molecule, yet the resultingnetwork is a cohesive and stable hydrogel. Therefore, when used as acartilage substitute in vivo, because a majority of the HA molecules inthe invented T-HA network or gel are essentially unaltered compared toHA in normal cartilage, it is believed that the body's native metabolicpathways (aided or unaided by cells provided within the T-HA network)may recognize the invented network as native biologic material, and willbe able to carry out ordinary synthesis and metabolism functions withrespect thereto. In addition, it is noted that HA is a highly ubiquitousmaterial in the body, and is non-immunogenic in humans. As a result, thecross-linked macromolecular network, comprised a majority of unalterednative HA, will have substantial application in a wide variety of tissueengineering applications where it is desirable or necessary to providesynthetic tissue in a human body. This represents a significant advanceover the state of the art. Therefore, quite surprisingly, a high degreeof tyramine substitution, e.g. greater than about 10-20%, may beundesirable; the above described experiments demonstrated that such highdegrees of substitution are unnecessary to provide a suitable T-HAnetwork. Preferably, a dihydroxyphenyl (e.g. dityramine) cross-linkedpolycarboxylate (e.g. HA) network has a hydroxyphenyl (tyramine)substitution rate of less than 50, preferably less than 40, preferablyless than 30, preferably less than 20, preferably less than 15,preferably less than 10, preferably less than 9, preferably less than 8,preferably less than 7, preferably less than 6, preferably less than 5,percent based on the total quantity of CO₂H groups present on thepolycarboxylate (HA) molecules.

Example 3

Conventionally, it has been believed that natural cartilage exhibits itsviscoelastic properties and its ability to resist deformation and absorbcompressive loads principally as a result of the repulsive forcesbetween negatively charged SO₄ ²⁻ groups on adjacent chondroitin sulfatechains present in the aggrecan matrix. An experiment was performed todetermine the efficacy of various macromolecular networks within thescope of the invention to resist deformation and absorb compressioncompared to natural cartilage. In particular, three such networks wereprepared, respectively, composed of the following: 1) dityraminecross-linked HA molecules (T-HA); 2) dityramine cross-linked chondroitinsulfate molecules in the form of aggrecan (T-Aggrecan); and 3) acomposite material composed of 50% T-HA and 50% T-Aggrecan. Formulationsof uncross-linked T-HA and T-Aggrecan were prepared and purified as inExample 1, each having a tyramine substitution rate of about 5%. Fromthese T-HA and T-Aggrecan formulations, five different concentrations ofthe T-HA alone, T-Aggrecan alone and a 50:50 mixture of T-HA andT-Aggrecan were prepared:

Concentration 1: 6.25 mg total T-GAG/mL water

Concentration 2: 12.5 mg total T-GAG/mL water

Concentration 3: 25 mg total T-GAG/mL water

Concentration 4: 50 mg total T-GAG/mL water

Concentration 5: 100 mg total T-GAG/mL water.

The notation T-GAG is used herein to embrace both T-HA and T-Aggrecan.Though aggrecan technically is not a glycosaminoglycan (GAG), forpurposes of this example T-GAG nonetheless is defined to embrace bothT-HA and T-Aggrecan hydrogels. Each of the above preparations was thenreacted in the presence of hydrogen peroxide and horseradish peroxidase,also as in Example 1, to form dityramine cross-links between the T-GAGmolecules and provide respectively Hydrogels 1, 2, 3, 4 and 5 for eachof the three material compositions. Each of the fifteen hydrogels (fiveconcentrations for each of the three material compositions) was found tobe a stable and substantially coherent material with the physicalproperties of each hydrogel varying relative to the concentration ofT-GAG in the preparation from which it was made. For example,qualitatively T-HA Concentration 1 resulted in T-HA Hydrogel 1 havingrigidity and rheological properties comparable to that of Vaseline orjelly; the hydrogel was stable and coherent yet could be caused to flowor spread on application of an external force, e.g. from a spatula orother conventional tool. T-HA Hydrogel 1 exhibited excellent adhesiveproperties making it an ideal candidate for a nonallergenic coatingmaterial for surgical instruments during surgery, e.g. ophthalmologicsurgery. T-HA Hydrogel 2 was more rigid than T-HA Hydrogel 1 due to thegreater concentration of T-HA in the preparation from which it was made,and the consequent predicted decrease in intramolecular cross-linkingand increase in intermolecular cross-linking associated with increasedT-HA concentration. T-HA Hydrogel 2 exhibited rheological and rigidityproperties characteristic of gelatins, with a degree of viscoelasticreboundability on external loading. On greater loading, T-HA Hydrogel 2was found to break up into smaller pieces instead of flowing, alsocharacteristic of a gelatinous material. T-HA Hydrogel 3 had theproperties and consistency of a dough or malleable paste, also notflowing on application of an external loading force. This material alsoexhibited substantially greater viscoelastic properties compared to T-HAHydrogels 1 and 2. T-HA Hydrogel 4 was a highly rigid and coherent gelthat strongly resisted breaking up on application of an external loadingforce. T-HA Hydrogel 4 was a highly resilient rubber-like compositionthat actually generated substantial springing force upon suddencompression (e.g. dropping onto the floor). This ability of T-HAHydrogel 4 to generate such a springing force in response to a suddencompression may make this material ideal for certain jointreplacement/repair applications where the joint undergoes repeated andperiodic compressional loading (e.g. the ankle joint). In addition tothe properties described for T-HA Hydrogel 4, T-HA Hydrogel 5 hadcartilage-like properties with both the appearance of articularcartilage and the feel of cartilage upon cutting with a surgical blade.

Confined compression tests were performed to quantitatively determinethe compressive mechanical properties of the fifteen differenthydrogels. A custom built polycarbonate confining chamber, and porouspolypropylene filter platen (20 μm pores, 20% porosity) were used toperform the confined compression testing. Five cylindrical plugs (7.1 mmin diameter, approximately 3 mm in thickness) at each hydrogelconcentration for each of the three material compositions were madeusing the confining chamber and the freeze-thaw technique described inExample 4 below. The following testing protocol was followed for aseries of stress relaxation tests in confined compression. All testingwas performed using an Instron 5543 machine under computer control,which recorded the time-displacement-load data at a frequency of 10 Hz.A ±5 N or ±50 N load cell (Sensotec) was used to monitor load throughouteach test. A step of 30 μm (30 μm/sec), representing 1% strain, wasapplied until the sample reached equilibrium. This was defined as arelaxation rate that slowed to less than 10 mN min⁻¹, at which time thenext step was automatically started, until 20 cycles (representingapproximately 20% strain) were completed. The thickness of each sampletested in confined compression was determined mechanically, by measuringthe displacement at which the compressive response initiated relative tothe bottom of the chamber as measured with the Instron 5543 machine. Themeasured thickness was used to calculate the strain percentage for eachstep.

The compressive mechanical properties of the fifteen hydrogels weredetermined as described in the preceding paragraph. Load data wasnormalized by sample cross-sectional area (39.6 mm²) to compute stress.The equilibrium stress was plotted against the applied strain for eachmaterial formulation. The aggregate modulus at each step was defined asthe equilibrium stress divided by the applied strain. For each material,the aggregate modulus was defined as the slope of the equilibriumstress-strain data in the most linear range. FIGS. 4 a, 4 b and 4 cdisplay the equilibrium compression behavior for the five concentrationsof T-HA, T-Aggrecan and 50:50 T-HA/T-Aggrecan composite hydrogels,respectively. All fifteen hydrogels were testable in confinedcompression, and demonstrated characteristic stress relaxation responsestypical of biphasic materials (such as cartilage). The aggregate modulifor the 6.25 mg/ml and 12.5 mg/ml T-GAG hydrogels were 1-2 orders ofmagnitude lower than articular cartilage. The 25 mg/ml T-GAG hydrogels,as well as the 50 mg/ml T-aggrecan hydrogel, displayed aggregate modulion the order of, but at least 30% lower than that of articularcartilage. All the 100 mg/ml T-GAG hydrogels, as well as the 50 mg/mlT-HA and the all the composite hydrogels, displayed aggregate moduli,equal to or exceeding reported literature values for articularcartilage. These data demonstrate the ability to characterize hydrogelsusing standard mechanical assays, and to generate hydrogels with similarmechanical properties to a wide variety of tissues including that ofarticular cartilage, using a variety of glycosaminoglycans as thehydrogel scaffold material.

The aggregate moduli for the five concentrations of the T-HA, T-Aggrecanand composite materials composed of 50% T-HA and 50% T-aggrecan aresummarized below in Table 1.

TABLE 1 Aggregate Modulus (MPa) HA Aggrecan 50/50 Composite (n = 5) (n =5) (n = 5) 6.25 mg/ml *0.024 ± 0.014  *0.008 ± 0.003  0.064 ± 0.019 12.5mg/ml 0.072 ± 0.024 *0.032 ± 0.006  0.108 ± 0.004 25 mg/ml 0.482 ± 0.1310.111 ± 0.021 0.277 ± 0.021 50 mg/ml 1.023 ± 0.164 0.366 ± 0.065 0.754 ±0.071 100 mg/ml 1.241 ± 0.351 0.748 ± 0.179 2.850 ± 0.377 *(n = 3)

FIG. 4 d shows the measured aggregate modulus as a function ofconcentration for the T-HA, T-Aggrecan and composite hydrogels. As theconcentration of the T-HA hydrogels increases, a plateau is reached forthe aggregate modulus while the T-Aggrecan hydrogels display a linearrelationship. Interestingly, the composite hydrogels show a relationshipindicative of an exponential increase in compressive properties asconcentration increases. This indicates that the moduli of otherhydrogel materials can be predicted by further exploring and modelingthese relationships.

Based on the above experiments it was surprisingly and unexpectedlydiscovered that a dityramine cross-linked GAG network (HA or aggrecan)will produce a coherent hydrogel material whose rigidity and otherphysical (Theological) properties can be tuned by varying the T-GAGconcentration prior to cross-linking the tyramine groups to suit aparticular application. The coherence and elastic properties of thesehydrogels was observed even absent any (or substantially any) SO₄ ²⁻groups in the network to supply the charge-to-charge repulsive forces togenerate the material's compression resistance and elasticity. This wasa highly surprising and unexpected result with substantial positiveconsequences in tissue engineering applications. Hyaluronan is a highlyubiquitous and non-immunogenic molecule found in humans. Therefore,hydrogels comprised of dityramine cross-linked hyaluronan networks canbe used to provide suitable tissue replacement materials that can beimplanted within a human body, whose rigidity can be tuned based on theapplication as evidenced by this example. As these materials can or willbe composed of predominantly unaltered hyaluronan which isnon-immunogenic, the hydrogels should result in zero or substantiallyzero immune response. This is an important advantage over manyconventional tissue engineered materials whose formation chemistriesprevent their application in vivo due to harsh reaction conditions orreagents, and whose final chemical structures are more likely to inducean immune response.

Example 4

A number of methods of preparing hydrogels such as those described inExample 3 have been developed to cast or form the hydrogel into apredetermined three-dimensional shape. This is important for myriadtissue engineering applications where it is necessary to provideartificial tissue material to fill a native tissue defect or void withina patient. A first method is to employ an in situ forming techniquewhere the hydrogel is formed in place, i.e. in position and in the shapeof its final application and structure. The in situ formation method hasbeen carried out experimentally as follows. Tyramine-substitutedhyaluronan (T-HA) was prepared via the carbodiimide-mediated pathwaydescribed herein. Following dialyzation to remove unreacted tyramine,EDC, NHS, etc., and dissolution at the desired concentration in PBS (seeExample 1 above), a small quantity of horseradish peroxidase enzyme wasadded to the T-HA liquid preparation to form a first solution. Thisfirst solution was provided into a laboratory container (to simulate anin vivo situs) having a specific interior geometry. Subsequently, asecond solution was prepared containing very dilute hydrogen peroxide(0.012%-0.00012% final concentration). A small volume of this secondsolution relative to the first solution was then injected into thecontainer already containing the first solution to initiate thedityramine cross-linking reaction to yield the hydrogel. Hydrogelsprepared by this technique have been prepared having varying rigidityand rheological properties as described above in Example 3, andconformed well to the interior surface contour of the container in whichthey were formed. Because the principal reagents (H₂O₂, hyaluronan andperoxidase) are either nonallergenic or diffusible molecules, andbecause the cross-linking reaction proceeds under metabolic conditionsof temperature and pH, this technique can be performed in vivo at asurgical situs in a patient as a surgical procedure to produce adefect-conforming hydrogel. This method is particularly attractive forreconstructive facial surgery in which the uncross-linked T-HApreparation (with peroxidase) can be injected and manipulatedsubcutaneously by the surgeon to produce the desired facial contours andthen the hydrogel subsequently cross-linked by injection of a smallvolume of the hydrogen peroxide solution.

A second method is a porous mold technique and is suitable for forminghydrogels into more complex three-dimensional structures. In thistechnique a porous hollow mold is first cast conforming to the shape andcontour of the intended final structure. For illustration, a mold can beprepared having an interior surface in a cuboid shape if a cuboid shapedhydrogel were desired. The mold can be prepared or cast via conventionaltechniques from conventional porous materials, e.g. plaster of paris,porous or sintered plastics or metals, etc. In a particularly preferredembodiment the mold is prepared using a cellulosic dialysis membrane.The first and second solutions are prepared as above, and the firstsolution is provided into the hollow mold cavity of the porous mold.Subsequently, the now-filled mold is submersed in a bath of very diluteperoxide. The macromolecular T-HA and peroxidase molecules are unable todiffuse out of the porous mold due to their size, however the very smallperoxide molecules (H₂O₂) are able to diffuse in and react in thepresence of the peroxidase enzyme to yield dityramine cross-links. It isinherent in this method that cross-linking occur from the outside inwardto produce the finished hydrogel shape, and a certain degree of trialand error may be required to determine optimal or sufficient immersiontimes in the peroxide bath. Determination of these time periods iswithin the skill of a person having ordinary skill in the art.Successfully completed three-dimensional hydrogel shapes have beenprepared in laboratory bench experiments via this porous mold technique.

A third method is a freeze-thaw technique that is suitable for castinghydrogels according to the invention in highly intricate predeterminedthree-dimensional shapes, e.g. having internal folds such as a humanear. In this technique, a mold is prepared from a soft or malleablematerial such as a polymeric material having a low glass transitiontemperature, e.g. below −80° C. The preferred mold materials aresilicones having low glass transition temperatures, such aspolydimethylsiloxane whose glass transition temperature is about −127°C., however other suitably low glass transition (e.g. below −80° C.)silicones, as well as other polymers, can be used. The silicone(preferred material) is first prepared such that it has an inner moldcavity conforming to the surface shape, contour and volume of a desiredhydrogel part via any conventional or suitable technique (i.e.press-molding, carving, etc.). First and second solutions are preparedas above, and the first solution is provided into the inner mold cavityof the silicone mold. The now-filled silicone mold is then cooled toabout −80° C. by contacting with solid CO₂ (dry ice). Because the firstsolution is principally water, it freezes into a solid ice formconforming to the shape and contour of the inner mold surface. However,the silicone mold, having a glass transition temperature below −80° C.,remains soft and malleable and the solid ice form of the first solutionis easily removed. Because the first solution expands as it freezes,suitable mechanical hardware should be used to ensure the silicone molddoes not deform or expand as the solution freezes. Preferably, portholes are provided in the mold to allow for expansion and discharge ofthe first solution as it expands during the freezing process.

Once the solid ice form of the first solution has been demolded, minutedefects or flaws in the three-dimensional structure can be repaired bycarving with a suitable tool, and more of the liquid first solution canbe added to fill surface voids, which liquid instantly freezes oncontact with the solid ice form. Also, the ice form can be placed backon the dry ice surface if desired to ensure uniform temperature andfreezing of any added first solution material. Once thethree-dimensional shape of the ice form has been perfected, it isimmersed in a liquid peroxide solution to initiate thawing of the frozenwater and dityramine cross-linking from the outside-in. This is possibledo to the rapid kinetics of the cross-linking reaction. Cross-linking isdetermined to be complete once the last remaining frozen water hasmelted at the center of the forming hydrogel form, which can be easilyobserved because the forming hydrogel is substantially clear.

Very successful experiments have been performed according to thisfreeze-thaw technique to produce a solid hydrogel in the shape of ahuman ear. Other structures that could be formed by this method, such asintervertebral discs, meniscus, etc. will be evident to those skilled inthe art. It should be noted in this freeze-thaw technique, the thresholdglass transition temperature of −80° C. for the mold material isselected to correspond roughly with the surface temperature of solid CO₂(dry ice), to ensure the mold material does not become brittle when thefirst solution is frozen to produce the solid ice form. However, ifanother cooling material, other than CO₂ is used, then the thresholdglass transition temperature for suitable mold materials may be adjustedaccordingly.

For the three methods of hydrogel formation described above, the firstsolution contained both the peroxidase and T-HA, while the secondsolution contained the peroxide. While it may be possible to switch theperoxidase and peroxide in the first and second solutions respectively,it is less preferred to provide the peroxide in the first solution withthe T-HA. This is because once the peroxide, peroxidase and T-HA arecombined, the T-HA rapidly begins to form a cross-linked macromolecularnetwork. If the peroxidase (which is a macromolecular molecule) is notalready uniformly distributed with the T-HA it may be unable orsubstantially hindered from diffusing through the pore structure of theforming hydrogel to facilitate uniform cross-linking throughout theentire T-HA/peroxide solution. The result could be non-uniform and/orincomplete cross-linking of the T-HA and a non-uniform hydrogel.Conversely, the relatively small peroxide molecule (hydrogen peroxide isonly one oxygen atom larger than water) can diffuse through the hydrogelpore structure with relative ease, resulting in a uniform hydrogelstructure.

In addition, the macromolecular size of the peroxidase allows it to besimilarly retained as the T-HA within porous molds that are only porousto small molecular weight peroxides which easily and uniformly diffusethrough both the molds and newly forming macromolecular networks (i.e.hydrogels). For these reasons it is preferred to start with theperoxidase uniformly distributed with the T-HA in the first solution,and to provide the peroxide separately in the second solution.

A fourth method is an alternating sprayed or brushed layering technique.The first solution is prepared as described above and contains both theperoxidase and T-HA. However, the second solution not only contains theperoxide as described above, but also T-HA at the same concentration asin the first solution. Then a thin layer of the first solution isapplied at the desired location (in situ) followed by an overlying thinlayer of the second solution. This procedure is repeated such thatalternating layers of the first and second solutions are successivelyapplied until the defect or application situs has been completed. Thevery thin alternating layers of the first and second solutions promotevirtually complete dityramine cross-linking ensuring a highly coherentfinal hydrogel having the desired rheological properties based on theinitial T-HA concentration of the two solutions. The thin nature of thelayers is desirable to ensure that free radicals produced by theperoxidase in the first solution layers are able to penetrate completelyadjacent second solution layers and complete cross-linking independentof peroxidase diffusion into the second solution layer (see above). T-HAis included in both solutions to ensure uniform T-HA concentrationthroughout the final hydrogel. This technique has been performed inlaboratory bench experiments and has provided contour-conforming andvolume-filling coherent hydrogels. This technique is highly applicablewhere it is desired to provide a thin, but variable layer of tyraminecross-linked HA, such as on the surface of a denuded osteoarthriticjoint in which little if any native healthy cartilage remains in thepatient at the implant site.

All four of the above techniques have been described with respect todityramine cross-linked hyaluronan, however it will be understood thatother combinations within the scope of the present invention (otherdihydroxyphenyl cross-linked macromolecules, such as polycarboxylates,polyamines, polyhydroxyphenyl molecules and copolymers thereof) can bemolded via the above techniques.

Example 5

Rat chondrocytes were embedded in (cross-linked) T-HA hydrogels tomeasure their ability to survive the cross-linking reaction. Isolatedchondrocytes were suspended in the 1.7% and 4.7% T-HA hydrogelsdescribed in Example 2 by providing these live cells to the firstsolution to be co-dispersed with the T-HA and peroxidase, followed byintroduction of the peroxide-containing second solution to initiatedityramine cross-linking. The chondrocyte-embedded 1.7% and 4.7% T-HAhydrogels exhibited uniformly distributed chondrocytes with the opticalclarity of the gels allowing visualization throughout the gel. Glucoseutilization was used as an indicator of cell viability aftercross-linking to form the hydrogels as chondrocytes are voracious withrespect to glucose consumption, depleting the medium of glucose in lessthan 24 hours. The results showed that chondrocytes embedded in T-HAhydrogels showed essentially the same glucose consumption profile over24 hours as the same chondrocytes cultured in monolayer (FIG. 5). Thiscontinued for up to 7 days indicating that the cells were alive andmetabolically active. Medium glucose was measured by standard hexokinaseassay.

Fluorescent images of frozen sections of T-HA hydrogels containing bothchondrocytes and cartilage tissue were also generated. HA samples fromboth the hydrogel scaffold and cartilage matrix were visualized byfluorescent staining with biotinylated HA binding protein (b-HABP)reagent while cell nuclei were visualized with standard DAPI stain. Theb-HABP reagent is prepared from purified cartilage aggrecan (the G1domain only) and link protein, and recognizes and irreversibly binds tostretches of native HA equivalent to those normally bound by aggrecanand link protein in cartilage. The results showed a more intensestaining of the T-HA hydrogel with b-HABP than the cartilage as thehyaluronan in the tissue is already occupied by native aggrecan and linkprotein. No visible distinction could be seen between the T-HA scaffoldof the hydrogel and the matrix of suspended cartilage tissue suggestingseamless integration. These results demonstrated the feasibility ofmaintaining the viability of chondrocytes during the hydrogelcross-linking reactions, and the ability of the hydrogel to integrateseamlessly into existing cartilage matrix, both of which areadvantageous for application to cartilage repair. The results alsodemonstrated that sufficient stretches of the T-HA remain chemicallyunaltered, and available for binding by newly synthesized aggrecan andlink protein in situ. The results also demonstrated that oxygen, carbondioxide, glucose and insulin are diffusible through T-HA hydrogelsaccording to the invention at a rate that is not limiting to chondrocytemetabolism, which is important not only to the development of cartilagesubstitutes but to other applications such as glucose sensor design anddevelopment of an artificial kidney.

In order to include cells such as chondrocytes in hydrogels molded intointricate anatomical shapes using the freeze/thaw technique described inExample 4, it is desirable that the enzyme driven cross-linking reactionproceed in the presence of standard cell freezing solutions such asthose containing 10% dimethylsulfoxide (DMSO)/90% fetal bovine serum(FBS). This has been demonstrated in the laboratory for all of the T-HAhydrogel formulations described in Example 3. The ability to directlyincorporate a solution containing 90% FBS also demonstrates the abilityto include bioactive factors such as growth factors, hormones andfactors controlling cell differentiation, as these are normal componentsof FBS.

Example 6

An experiment was conducted whereby a T-HA hydrogel as describedhereinabove was implanted into Yucatan minipigs in order to repairarticular cartilage defects. Following is a description of thatexperiment, including the experimental methods and results obtained,after a brief discussion of the background for this application.

Background Tissue Description

Articular Cartilage Structure and Function—As discussed above, articularcartilage is the resilient load-bearing tissue that forms thearticulating surfaces of diarthrodial joints. It absorbs mechanicalshock and deflects or spreads applied load over greater surface area ofsubchondral bone. It consists primarily of a large extracellular matrix(ECM) with a sparse population of highly specialized cells(chondrocytes) distributed throughout the tissue. The primary componentsof the ECM are water, cartilage aggregates and type II collagen.Cartilage aggregates are composed of hyaluronan (HA), aggrecan (thelarge cartilage-specific proteoglycan), and link protein (LP), a smallglycoprotein. Aggrecan contains a central core protein to which isattached ˜100 chondroitin sulfate (CS) chains. The core protein hasthree globular domains with the N-terminal globular 1 (G1) domain havingbinding sites for both HA and LP. LP has sequence homology to the G1domain of aggrecan, and contains binding sites for both HA and the G1domain of aggrecan. Each cartilage aggregate is composed of a single HAchain, to which are attached hundreds of aggrecan/LP duplexes. Theselarge cartilage aggregates are trapped at one fifth of their freesolution volume within a tight meshwork of type II collagen fibers,which resist further swelling. This molecular architecture contributesto the tissues mechanical properties and function as described below.

Swelling Pressure—The HA and CS chains in cartilage aggregates containrepeating carboxyl and/or sulfate groups. In solution, these groupsbecome ionized (COO⁻ and SO₃ ⁻), and in the physiologic environment theyrequire positive counter ions such as Na⁺ to maintain overallelectroneutrality. These free-floating ions within the interstitialwater are present at a higher concentration than that found in thesurrounding fluids (i.e. synovial fluid) giving rise to an osmoticpressure (Donnan pressure). In cartilage, ions are prevented fromflowing out of the tissue along the concentration gradient by the fixednature of their negative counter ions (i.e. the COO⁻ and SO₃ ⁻ groups onthe HA and/or CS chains), and the need to maintain electroneutrality.Water flow into the tissue to equilibrate the concentration gradient isresisted by the inextensible nature of the collagen meshwork preventingfurther swelling.

Alternatively, tight cartilage aggregate packing causes thefixed-negative charge groups to be spaced only 10 to 15 angstroms apart,resulting in strong charge-to-charge repulsive forces (electrorepulsiveforces). As with the Donnan effect, the tendency to swell to lessenthese repulsive forces is resisted by the inextensible nature of thecollagen meshwork. When compressed, the distances between charge groupsdecrease, thus increasing the charge-to-charge repulsive forces andincreasing the free-floating positive counter ions concentration. Thusboth the Donnan and electrorepulsion effects are intensified bycompression. Both effects contribute to the swelling pressure ofarticular cartilage and its ability to resist deformation and absorbcompressive loads.

Stress Shielding Effect—Articular cartilage is often described as aviscoelastic, biphasic material, composed of a solid phase (cartilageaggregates, collagen, etc.) and a fluid phase (water and dissolvedions). The macromolecular architecture of the ECM of articular cartilagefunctions to deflect applied forces during loading from the wearsusceptible solid phase of the tissue to the wear resistant fluid phaseor water. This stress shielding occurs due to the elegant design of thecartilage ECM which produces a material with very low permeabilitycreating a drag during interstitial fluid flow. Interstitial fluidpressure is generated during compressive loading, and during dynamicloading, is the primary force responsible for supporting the appliedload with matrix compression a minor factor. During compression, theporosity is reduced further, which increases the already high frictionaldrag forces. The load support is gradually transferred from the fluidphase (as the fluid pressure dissipates) to the solid phase. Typically,for normal cartilage, this equilibration process takes 2.5 to 6.0 hoursto achieve. Thus, load support through fluid pressurization predominateswithin the tissue.

Need for Synthetic Material

Increased water content and decreased proteoglycan content are the mostapparent early changes in osteoarthritic cartilage. These changesreflect an increase in tissue permeability. Increased permeabilitydiminishes the fluid pressurization mechanism of load support incartilage (stress shielding), requiring the collagen-aggrecan solidmatrix to bear more load, which may be an important contributing factorin the development and progression of cartilage degeneration.Bioartificial cartilage substitutes that do not mimic the lowpermeability of normal, healthy articular cartilage may be predisposedto degeneration by a similar mechanism.

One of the most difficult challenges facing orthopedic surgeons istreating patients who have suffered focal cartilage lesions, but are tooyoung or too active for a total joint replacement. These localizedcartilage defects can be very debilitating. Restoring these localizedareas without a total joint replacement would be a preferred approachwith significant benefits including reduced surgical requirements,shorter recovery times, lower cost, and slowing or arresting the furtherdegradation of the load bearing surface.

This example demonstrates the application of a tyramine-substituted HA(T-HA) hydrogel for repair of this type of localized cartilage defect.

Experimental Description

Design of Extracellular Matrix Material having Desired Properties

Natural articular cartilage has the elastic as well as physical andchemical properties described above, which impart its unique ability toabsorb mechanical loads and to deflect impact loads away from thesubchondral bone. To produce a suitable synthetic cartilage materialmade from a T-HA hydrogel as disclosed herein, it was important todesign the macromolecular network for the hydrogel so as to emulatethose properties as nearly as possible through judicious selection ofreagent concentrations, cross-linking conditions, incorporated livingcells as well as other molecules, etc.

The results from confined compression testing of T-HA hydrogels (seeExample 3 above) provided an initial basis for the synthesis of anappropriate synthetic T-HA material having properties matched to thosemeasured for normal articular cartilage. They also illustrate thespectrum of material properties that can be manufactured from a singleformulation of T-HA. Based on those data, a T-HA hydrogel composed of,inter alia a macromolecular network of dityramine cross-linkedhyaluronan molecules was selected based on the following criteria inorder to produce a synthetic implantable cartilage material.

A material composed solely of HA was chosen because hydrogelcompositions with the compressive properties of cartilage could beformed using HA alone as the scaffold material (Example 3), whileavoiding possible host response to the protein component of aggrecan ifit were used as a scaffold material. Reaction conditions (tyramine/EDCratio) were chosen to produce a percent tyramine substitution of 5%(Example 2) as this provided sufficient cross-linking to produce amaterial with the compressive properties of cartilage (Example 3) whilemaintaining the majority of the native HA structure. HA was substitutedwith ˜5% tyramine, as described in Example 1, except that the HA wasdissolved at 5 mg/ml rather than 1 mg/ml to conserve reagents. Theabsolute concentration of all other reagents remained the same so thatthe tyramine and EDC were at a 2-fold rather than 10-fold molar excessbased on the molar concentration of HA carboxyl groups. A concentrationof 125 mg/ml of HA in sterile saline was chosen as this concentration insaline had a compressive aggregate modulus most closely resembling thatof articular cartilage (saline data not shown). It was also theconcentration deemed most appropriate based on the experience of ourclinician collaborator. Peroxidase was added at 10 U/ml prior toapplication in an in situ cross-linking protocol as described below. Thein situ cross-linking protocol was chosen as it provided the bestopportunity for integration with surrounding cartilage matrix. It alsoallowed easy and complete filling of the surgically produced cartilagedefect without the need to know or measure the defect's exactdimensions. Pre-cast (in vitro cross-linked) plugs would either requireexact dimensions or sculpting of pre-formed shapes to fit the defect. Nocells or bioactive factors where added in this experiment as thisexperiment was intended to evaluate the hydrogel material independent ofcomplicating factors derived from inclusion of cells or bioactivefactors. However, cells or bioactive factors could be included asdescribed hereinabove to produce desired effects.

Surgical Procedure

Pre-Operative—After arrival in the biological resource unit, Yucatanminipigs (˜7-8 months of age, ˜30-35 kg) were maintained for a minimumof 7 days to ensure full acclimatization. After pre-medication withKetamine (20 mg/kg I.M.) as anesthesia and Ambipen (40,000 U/kg I.M.) asprophylaxis antibiotic, the animal's rear legs were shaved and paintedin Betadyne as simultaneous bilateral surgery of both knees wasperformed. A general anesthesia was maintained by inhalation withIsoflurane (1-2.5% volume) in O₂ following intubation. Thiopental wasused as needed (to effect 25 mg/ml I.V.). During surgery, the animal wasmonitored for heart rate, respiration rate, body temperature, etc.

Opening—A longitudinal midline skin incision was made and carried downsharply through the pre-patellar bursa. Electrocautery was used forhemostasis. The lateral border of the patella was identified and alateral para-patellar arthrotomy was performed. The lateral retinaculumand musculature were tagged with #1 vicryl suture. The patella wasdislocated medially to expose the femoral trochlea.

Cartilage repair—As seen in FIG. 6, two circular full thickness chondraldefects (˜4.5 mm in diameter, panel B of FIG. 6) were created in themedial trochlear facet of the femoral chondyle (panel A of FIG. 6) usingan Acufex 4.5 mm mosaiaplasty chisel and a sharp, curved currette takingcare as much as possible not to disrupt the osteochondral plate. Thedefects were filled with in situ cross-linked T-HA hydrogel (125 mg/mlin sterile saline) as follows to produce a hydrogel implant having thecomposition described above in order to reproduce the in vitro measuredcompressive properties of natural cartilage as described above.Initially, each defect was rinsed with 0.01 cc of 0.6% hydrogenperoxide, and then immediately blotted dry with sterile gauze.Subsequently, a plug containing 0.15 cc of uncross-linked hydrogel paste(panel C of FIG. 6) having the composition and prepared as describedabove was inserted into and used to fill each defect with the surgeonsmoothing the surface of the hydrogel implant with fingertips to matchthe contour of the articular surface. A sterile piece of filter paper(Whatman 50) soaked in 0.6% hydrogen peroxide was pressed against thesurface of the hydrogel implants for five minutes to cross-link thehydrogel in the defect. During the 5 minutes, the filter paper wasrubbed back and forth across the implant surface to prevent integrationwith the filter paper, and to effectively polish the implant surface.After 5 minutes the filter paper was removed, excess hydrogel trimmedfrom the site, and then ˜0.01 cc (1 drop) of 0.6% hydrogen peroxideadded to the surface of each implant plug (panel D of FIG. 6). ThePatella was reduced anatomically over the femoral trochlea. The Patellawas dislocated and reduced again to ensure secure primary stability ofthe hydrogel.

Closing—The joint was irrigated with sterile saline. The wounds wereclosed in layers with vicryl sutures. Specifically, the arthrotomy wasclosed with interrupted #1 vicryl suture, the subcutaneous tissue wasclosed with interrupted 2-0 vicryl suture and the skin layers wereclosed with interrupted 3-0 vicryl suture. No restriction of movementwas required after surgery.

Post-Operative—The animal returned to full weight bearing immediatelyfollowing surgery. Analgesia was provided by Buprenorphine (0.02 mg/kgI.M.) for 24 hours and a Fentanyl Patch (50 mcg/hr) for 3 post-operativedays. Post-operative prophylactic antibiotic in the form of cephalexin500 mg twice per day was given for 7 days. The animals were kept in aconventional animal run.

Post-Implantation Data—At one month post-implantation, the animal waseuthanized with overdose of the barbiturate, Beuthanasia D Special (1ml/10 kg B.W., I.V.) under general anesthesia. After euthanasia, theentire knee joint was carefully dissected, macroscopically evaluated andphoto documented. As seen in FIG. 7, macroscopic inspection of the kneesat 1 month revealed no significant effusion and no evidence ofinflammatory reaction. The lesions were partially filled with a whitematerial (the implanted T-HA hydrogel as well as other factors or cellswhich may have migrated into the hydrogel post-operatively) and thesurrounding articular cartilage and opposing articular surface (patella)were normal in appearance except for a slight abrasion appearing on theopposing articular surface as seen in panel B of FIG. 7. It is notevident this abrasion was the result of rubbing against the implants,particularly given its location on the Patella in a position which doesnot appear as though it would have abraded against the implants duringnormal articulation of the joint.

The results indicate no apparent negative effect on joint health as aresult of the hydrogen peroxide or peroxidase reaction used for in situcross-linking of the hydrogel, and demonstrated the utility of thehydrogels disclosed herein, comprising a dityramine cross-linkedhyaluronan macromolecular network, as a synthetic implantableextracellular matrix for use as a synthetic in vivo cartilagereplacement or implant material.

Example 7

An experiment was conducted whereby a T-HA hydrogel as describedhereinabove was implanted into canine and rabbit models in order torepair vocal cord defects as well as to augment vocal cords. Followingis a description of that experiment, including the experimental methodsand results obtained, after a brief discussion of the background forthis application.

Background Tissue Description

The vocal cords are complex, multilayered structures under very fineneuromuscular control. The overlying mucosa is composed of anon-keratinized, stratified squamous epithelium, with a multilayered,lamina propria deep to the epithelium. Underlying the lamina propria isa muscular layer consisting of the thyroarytenoid muscle which insertsinto the thyroid cartilage anteriorly and the vocal process of thearytenoid cartilage posteriorly. The thyroarytenoid muscle can stiffenor relax, altering the tension on the lamina propria and therebyaltering the vibratory dynamics of the epithelium, which produces thefinely coordinated vibrations responsible for high quality speechproduction.

The biomechanics of human voice production have been attributed to theaction of certain biological macromolecules naturally found within theextracellular matrix (ECM) of the lamina propria. Hyaluronan (HA) is aubiquitous molecule, which is most concentrated in specialized tissuessuch as the vocal cords, synovial fluid, umbilical cord, dermis, andcartilage. In these tissues, its function is manifold, influencingtissue viscosity, shock absorption, wound healing, and space filling.

The unique structure of HA elucidates its multiple functions. Itconsists of D-glucuronic acid and N-acetylglucosamine arranged inrepeating disaccharide chains, containing as many as 30,000 repeatingdisaccharide units with a mass of more than 10 megadaltons. As HA is apolysaccharide instead of a protein, it is non-antigenic. Underbiological conditions, it is a negatively charged, randomly coiledpolymer, filling a volume more than 1,000 times greater than expectedbased on molecular weight and composition alone. The strong negativecharges attract cations and water, allowing it to assume the form of astrongly hydrated gel, and giving HA its unique viscoelastic andshock-absorbing property.

Vocal cord viscoelasticity is essential to high quality voiceproduction, as it directly affects the initiation and maintenance ofphonation and the regulation of vocal cord fundamental frequency. HA inthe human glottis is concentrated in the lamina propria and itsimportance has been quantified by comparing the biomechanical propertiesof cadaveric vocal cords with and without HA. Treatment of the vocalcords with hyaluronidase led to a 35% average reduction in vocal cordstiffness and a 70% mean reduction in high frequency vocal foldviscosity, thus illustrating the significance of HA in these tissues.

Need for a Synthetic Material

Vocal Cord Repair—Defects in the vocal cords have a dramatic effect onvocal production. Arising either de novo, or resulting from surgicalintervention, heterogeneous masses within the vocal cords disrupt thefinely coordinated vibrations responsible for high quality speechproduction. Patients with de novo lesions usually present early in thedisease course, due to persistent hoarseness. When presenting with anearly stage malignant process (T1-T2 in the Tumor, Node, Metastasisstaging system), patients undergo rapid treatment consisting of eitherexternal beam radiation therapy or endoscopic surgery. Such patients arecounseled that poor post-treatment voice quality is an expected sideeffect of effective tumor eradication. When presenting with a presumedbenign process, patients are faced with a conundrum, for the surgicaltreatment often produces speech quality as poor as that caused by thelesion itself. Unfortunately, current standard laryngeal operativetechnique cannot provide for effective removal of either benign ormalignant lesions without causing poor vocal outcomes secondary to vocalfold scarring. This is due to the mechanism of wound healing in theunique anatomy of the larynx. The superficial, vibratory surfaces of thevocal cords become tethered to the deeper layers by the post-treatmentscar, preventing physiologic phonatory oscillation.

HA in the human glottis is concentrated in the lamina propria, ahistological layer separating the vocalis muscle from the overlyingepithelium. The lamina propria allows the epithelium to vibrate over thetaut vocalis muscle, like waves propagating over a pond. This “mucosalwave” is the sine qua non of effective speech production. In thepresence of benign or malignant lesions of the vocal cords, the mucosalwave is disrupted. Even in the normal process of healing, scar bands anddisorganized collagen “tether” the superficial mucosa to the deeperlayers of the vocal cords, disrupting the normal mucosal wave andimpairing vocal production. The shock absorbing nature of HA allows itto act as a tissue damper, protecting the mucosal surfaces from theoscillatory trauma experienced during phonation. HA also appears tofacilitate wound repair by minimizing fibrosis and scarring, therebyprotecting the vocal cord from the permanent damage resulting fromtrauma.

The development of a technique, permitting the restoration of a fullyvibratory phonatory surface on vocal cords undergoing laser or coldsurgical treatment, would enable a large population of patients withboth benign and malignant processes to undergo treatment of their tumorswith the expectation of unprecedented post-operative speech outcomes.

Vocal Cord Augmentation—A variety of disorders and diseases adverselyaffect glottic function, vocal quality and the ability to communicate.Approximately 7 million people in the United States suffer fromdysphonia or voice impairment, and those affected by vocal cordparesis/paralysis are a significant subset of this population. It isestimated that 1-4% of all cardiac and thyroid surgeries in the UnitedStates result in vocal cord paresis or complete paralysis due toinadvertent vagus nerve or recurrent laryngeal nerve injury duringsurgery.

Another condition affecting vocal cord function is unilateral vocal foldparalysis (UVP). In UVP the problem is malposition of an insensate vocalcord. While medialization results immediately following nerve injury dueto opposing tensions of the laryngeal adductors and abductors, theparalyzed vocal cord rapidly lateralizes to a paramedian position. Thearytenoid cartilage prolapses into the larynx following recurrentlaryngeal nerve injury, resulting in a change in vertical height of thevocal cord, as well as decreased dynamic tension often resulting invocal fold bowing. Atrophy with resultant shortening and bowing of thevocal cord occurs later as the thyroarytenoid muscle atrophies due to alack of neural stimulation. As a result of atrophy and lateralization,the contralateral vocal cord cannot fully contact the paralyzed cord,leading to manifestations of UVP.

Such manifestations include breathy hoarseness, a weak cough, aninability to valsalva (protect the airway), and difficulty swallowing;complications include aspiration (solids and liquids) and recurrentpneumonia. This can result in a life-threatening condition due toincreased incidence of recurrent pulmonary infections.

As only one functional vocal cord is required for normal voiceproduction, successful treatment consists of “medialization” of theparalyzed vocal fold, thereby enabling it to contact the contralateralmobile vocal fold. This normalizes voice production and preventsaspiration minimizing the risk of aspiration pneumonia. Vocal foldparalysis is currently treated in two ways: open trans-cervicalapproaches or trans-oral endoscopic vocal cord injection, also known asinjection laryngoplasty therapy (ILT). Ishiki-type I thyroplasty is themost commonly performed trans-cervical approach, where a “window” iscreated in the thyroid cartilage to allow placement of a silasticimplant into the body of the atrophic, paralyzed vocal fold, in effectpushing it into a more medial position. This procedure has a permanenteffect, although complications include implant migration, extrusion, orinfection.

In ILT, the paralyzed vocal cord is medialized by the endoscopicinjection of an exogenous substance. A wide variety of synthetic andbiologic materials are currently available as an injectant for treatmentof UVP, including: gelfoam, hydroxyapatite, autologous fat or facia,acellular cadaveric dermis (Cymetra®), collagen or Teflon®/Gortex®.Unfortunately, all have proven to be less than ideal with nonefulfilling all desired criteria for the ideal material for long-termvocal cord augmentation. Such limitations create the need for eitherre-injection or over-injection to account for the projected loss ofvolume.

A biocompatible, injectable material such as a T-HA hydrogel asdisclosed herein can be designed to mimic the rheological properties ofthe natural vocal cord tissue and persist indefinitely in vivo withoutmigration. The design, chemistry, and material properties of a T-HAhydrogel as described herein can be tuned to produce an injectablebioimplant that is uniquely suited to otolaryngology treatments such asILT through judicious selection of component concentrations andcross-linking methodologies as described hereinabove.

A suitable biocompatible, high-longevity synthetic material also isdesirable to treat pre-existing sulcus or scarring that can develop dueto trauma or develop spontaneously with aging. Such a material alsocould be used advantageously in place of saline as a diagnostic andsurgical aid prior to vocal cord surgery. Conventionally, saline isinjected within the HA matrix of the vocal cord lamina propria between alesion to be surgically removed and the underlying ligament. This isdone to: a) determine if the lesion involves the underlying ligament;and b) to make the surgery easier by increasing the distance between thelesion and the ligament (cold instruments), or providing a heat sink(laser). Ligament involvement complicates the surgery with penetrationof the ligament to be avoided if possible. This procedure could benefitfrom the incorporation of hepatocyte growth factors in the hydrogel,which is used to increase HA production in the lamina propria anddecrease collagen production associated with scarring.

Experimental Description

Design of Extracellular Matrix Material having Desired Properties

An ideal synthetic matrix or biomaterial for vocal cord augmentationwill have the following characteristics: 1) biocompatible, so there isno unfavorable immunologic response; 2) easily injectable to allow asurgeon to control the exact amount and location of injection through asmall needle; 3) readily available with minimal preparation for optimaltime efficiency and potential application to the outpatient officesetting; 4) possess the same or similar biomechanical properties to thevocal fold component being augmented to cause minimal alteration in thenatural function of the augmented structure; 5) resistant to resorptionor migration, so that the initial augmentation result is maintained; and6) easily removable in the event of revision surgery.

The T-HA hydrogels disclosed herein meet all six of these criteria, mostimportant being that its biomechanical properties can be tuned throughjudicious selection of reactant/synthesis parameters and GAG (e.g. HA)concentration to produce the necessary macromolecular network forproducing a hydrogel having desired viscoelastic and biomechanicalproperties. Specifically, results of both in vitro and in vivopreliminary studies have allowed favorable comparisons to be drawnagainst the above criteria. First, transplantation of T-HA hydrogelsinto various animal species including rat, rabbit, dog and pig all havedemonstrated little to no host immune response. Second, uncross-linkedT-HA hydrogels at the concentrations required for vocal cordaugmentation at the lamina propria or muscle level easily pass through a21 gauge needle. Third, medical-grade HA is readily available and hasbeen used for years in FDA-approved formulations such as Healon andRestylane. Furthermore, uncross-linked T-HA hydrogel and hydrogenperoxide (cross-linking agent) solutions can be readily pre-made, inoff-the-shelf formulations that require no preparation by the surgeon.Fourth, the T-HA hydrogels disclosed herein can be formulated to matchthe mechanical properties of the various tissues of the glottisincluding the lamina propria, thyroarytenoid muscle, and thyroid orarytenoid cartilages. Fifth, the unique cross-links and non-proteinnature of the T-HA hydrogels have demonstrated resistance to resorptionin in vivo experiments. This implies that the initial augmentationresult will be maintained. Finally, formation of a solid continuousimplant through the novel in situ cross-linking protocol made possiblethrough the non-immunogenic enzyme-driven cross-linking architecturedescribed herein should prevent migration and allow for easy locationand removal of the hydrogel implant should revision surgery be required.

With respect to the fourth point above, the results from confinedcompression testing of T-HA hydrogels (see Example 3 above) provided aninitial basis for the synthesis of an appropriate synthetic T-HAmaterial having properties matched to those of normal vocal cord tissue.Based on those data, a T-HA hydrogel composed of, inter alia amacromolecular network of dityramine cross-linked hyaluronan moleculeswas selected based on the following criteria in order to produce asynthetic implantable vocal cord repair or augmentation material.

Choices of scaffold material (HA alone), percent tyramine substitution(˜5%), protocol for tyramine substitution of HA (modified from Example1), and the non-incorporation of cells and bioactive factors were asdescribed in Example 6. A concentration range of between 2.5 and 10mg/ml of T-HA hydrogel in sterile saline most closely matched theTheological and vibratory properties of the vocal cord lamina propria.The 2.5 mg/ml concentration of T-HA hydrogel was deemed most appropriatebased on the extensive clinical experience of our cliniciancollaborators with both the vocal cord tissue and other injectablematerials used for vocal cord repair and augmentation. In vitrocross-linked hydrogel was used rather than an in situ cross-linkingprotocol based on the experience of our clinician collaborators withother injectable materials for vocal cord repair and augmentation. Invitro cross-linking was as described in Example 1. Based on the resultsof the rabbit and canine experiments described below a preferredembodiment for vocal cord augmentation using the disclosed hydrogelmaterial as envisioned by the inventors would use the in situcross-linking protocol and a T-HA concentration to more closely matchinjection into the deeper muscle layers of the vocal cord.

Surgical Procedure

Pre-Operative—After arrival in the biological resource unit, mongreldogs or New Zealand white rabbits (depending on the experiment,described below) were maintained for 7 days for full acclimatization.After pre-medication and general anesthesia per IACUC-approvedprotocols, each animal was intubated and maintained in stage IIIsurgical anesthesia. The animal was positioned supine. After graspingand superiorly retracting the tongue, a Dedo laryngoscope was placedtransorally providing good exposure of the larynx. The tip of thelaryngoscope was positioned several centimeters proximal to the superiorsurface of the true vocal cords. The laryngoscope was then suspended. Arigid videostroboscopic telescope was positioned above the true vocalcords, permitting complete inspection and imaging of the larynx.

Vocal Cord Repair—For vocal cord repair, lateral-based microflaps wereraised in both vocal cords of dogs, and then soft tissue defects werecreated equivalent to 50% of the vocal cord mass, including laminapropria and underlying muscle. One side underwent soft tissuereconstruction (filling) with the T-HA hydrogel, while the contralateralside served as an unrepaired (unfilled) control. The microflap was thenredraped over the hydrogel such that the epithelium was completelycontinuous. This study used 2.5 mg/ml ex vivo cross-linked T-HA hydrogelin saline (5% tyramine substituted) prepared as described above toapproximate the rheological properties of the lamina propria. Aftersurgery, the dogs were weaned from the anesthesia and transferred to therecovery room. The animals received analgesia for 1 to 2 days perIACUC-approved protocols.

Injection Laryngoplasty Therapy (Vocal Cord Augmentation)—Afteranesthesia, a 27 gauge laryngeal needle was used to inject approximately0.25 ml of 2.5 mg/ml ex vivo cross-linked T-HA hydrogel in saline (5%tyramine substituted) at the left anterior and posterior membranousvocal cord of rabbits. The injections were made in the superficial layerof the lamina propria.

Based on the results from the above experiments with dogs (vocal cordrepair) and rabbits (ILT), the following preferred embodiment for vocalcord augmentation using the disclosed hydrogel material is envisioned bythe inventors. For ILT, the lateralized and atrophied vocal cord areinjected with 50 mg/ml of the uncross-linked T-HA hydrogel in saline (5%tyramine substituted) with peroxidase at the level of the thyroarytenoidmuscle. Preferably one bolus of hydrogel would be used to obtain thedesired medialization, but no more than two. A 21 gauge laryngeal needleis used to inject the hydrogel. Cross-linking would be initiated byinjection of a small volume of dilute hydrogen peroxide through a 27gauge needle into the center of the implanted bolus of hydrogel usingthe 21 gauge needle, which would not have been withdrawn, for purpose oforientation. Cross-linking of the hydrogel into a solid implant would beachieved within minutes and verified by feel. After surgery, the animalswould be weaned from the anesthesia and transferred to the recoveryroom. The animals would receive analgesia for 1 to 2 days perIACUC-approved protocols. At the time of euthanasia, the vocal cords ofeach animal would be carefully dissected, macroscopically evaluated andphoto documented.

Post-Implantation Data

Vocal Cord Repair—At the time of euthanasia, the vocal cords of the dogswere carefully dissected, macroscopically evaluated and photodocumented. The degree of wound healing was assessed histologically by aconsulting pathologist with specific attention paid to inflammatoryinfiltrates, HA staining (normal matrix production), hydrogel stainingand collagen staining (scarring). FIG. 8 shows representativehistological results of control side (unfilled) and experimental side(T-HA hydrogel filled) vocal cords stained with alcian blue for one ofthe dogs three months following surgery. Gross observation indicated amore normal appearance and vibratory properties for the T-HAhydrogel-treated vocal cord compared to untreated controls. Thehistological results indicated significant scarring in the untreatedcontrol vocal cord along the wound track as indicated by a lack ofdeposition of GAG (i.e., HA) and increased collagen deposition whencompared to the T-HA-filled wound track of the experimentally repairedvocal cord.

Only small foci of T-HA hydrogel could be found in the experimentalvocal cord at 12 weeks, which show a minimal foreign body response witha layer of surrounding mast cells observed. This may indicatedegradation of the T-HA hydrogel with concomitant deposition of normalHA-containing tissue matrix. However, given the low concentration (2.5mg/ml) and thus the very fluid nature of the hydrogel used in thisstudy, it is more likely that much of the hydrogel was lost from thewound site prior to closure of the site as the epithelial microflapknitted to the opposed underlying tissue. Hydrogel between the microflapand opposed underlying tissue is actually predicted to inhibit theknitting process contributing to hydrogel loss from the wound site.Thus, the positive wound healing effect seen is believed due to only athin layer of hydrogel retained at the wound site rather than thevolume-filling bolus of hydrogel initially implanted. These resultsindicate the ability of the hydrogel to prevent scarring and match therheologic properties of the lamina propria.

Injection Laryngoplasty Therapy (Vocal Cord Augmentation)—At the time ofeuthanasia, the vocal cords of the rabbits were carefully dissected,macroscopically evaluated and photo documented. Histological evaluationby a consulting pathologist was used to assess the inflammatory responseand retention of the hydrogel. Representative histological results withAlcian Blue staining to detect the hydrogel and hematoxylin & eosin(H&E) for general morphology are shown in FIG. 9 for one of the rabbitsthat underwent the augmentation procedure at two weeks post-operatively.As seen in the figure, pockets of T-HA hydrogel could be found in theinjected vocal cords at 2 weeks, which show a minimal foreign bodyresponse with a layer of surrounding mast cells observed.

Example 8

An experiment was conducted whereby a T-HA hydrogel as describedhereinabove was implanted into a rabbit model in order to fill thevitreous cavity of the eye in order to prevent or treat vitreo-retinaldiseases such as retinal detachment. Following is a description of thatexperiment, including the experimental methods and results obtained,after a brief discussion of the background for this application.

Background Tissue Description and the Need for a Synthetic Material

Vitreo-retinal diseases, such as retinal detachment, diabeticretinopathy and others, are among the most common causes of blindness.The vitreous cavity of the eye normally is filled with a gel likesubstance. In retinal detachment surgery, the vitreous is surgicallyremoved (a procedure called “vitrectomy”), the retina is re-attachedagainst the back wall of the eye, and a replacement substance isinjected into the vitreous cavity. Vitreous substitutes are used for anumber of different purposes in the vitreous cavity of the eye. Theseinclude (1) achieving a long term tamponade after retinal re-attachmentsurgery to keep the retina apposed to the wall of the eye; (2) inintra-operative procedures such as unfolding of retinal tears, theremoval of subretinal fluid and the flotation and removal of dislocatedintraocular lens components; (3) for developing a sustained releasesystem that could maintain therapeutic drug levels in the posteriorsegment of the eye over long periods of time.

A number of different compounds are used as vitreous substitutes afterretinal re-attachment surgery. These compounds have physical propertiesthat permit successful retinal re-attachment but fail in other importantsurgical goals. Gases injected into the eye provide short term retinaltamponade but re-absorb quickly and cause significant optical distortionwhile they are in the eye. Perfluorocarbon liquids are effectiveintra-operative tools for flattening the detached retina but causeunacceptable toxicity when left in the eye for prolonged periods oftime. Silicon oil is used as a medium term retinal tamponade but alsocarries a risk of toxicity and causes significant optical distortion.

Alternatively, substitute vitreous compounds are desirable for use assafe, long term or time-released drug delivery vehicles in the eye. Manychronic inflammatory and infectious conditions of the eye, such assarcoidosis, idiopathic posterior uveitis and cytomegalovirus retinitis,necessitate intraocular injections of medication. Repeat intra-ocularinjections pose risks such as bleeding, retinal detachment andinfection. A stable, non-toxic vehicle is needed for sustainedintravitreal drug delivery.

Hyaluronan (HA) is an acellular substance that is an essential componentof natural vitreous in humans and other mammals. Formulations ofhyaluronan are already in use in some ophthalmic surgical procedures.For example, sodium hyaluronate is the most commonly used viscoelasticsurgical device for anterior segment and cataract surgery.Unfortunately, sodium hyaluronate and other previously tested hyaluronansubstitutes are dissolved relatively quickly in human tissues. Thesesubstances have not proven effective in vitreous surgery because oftheir failure to provide long-term retinal tamponade.

Experimental Description

Design of Extracellular Matrix Material having Desired Properties

The most common need for a vitreous replacement is during retinaldetachment surgery. A significant challenge is maintaining the retinaflat against the wall of the eye for a prolonged period of timepost-operatively. An ideal vitreous substitute should be optically clearto allow maximum visual rehabilitation during the recovery period.Finally, retinal detachments that occur inferiorly in the eye pose aparticular challenge. For the retina to remain flat post-operatively,the vitreous replacement must be directly apposed to the area of theretinal tear. To tamponade inferior breaks the patient must often lie ina face down position for weeks after the surgery. None of the vitreoussubstitutes in use today satisfy all of the current clinical needs.

There is a need for a non-toxic, optically clear, vitreous substitutethat will result in improved surgical results and post-operative visualrehabilitation in patients undergoing retinal detachment surgery.

Hydrogels made from a tyramine-substituted and cross-linked hyaluronan(T-HA) macromolecular network as disclosed herein present an idealchoice for a synthetic vitreous material. Specifically, the novelenzyme-driven cross-linking chemistry described above for cross-linkingthe tyramine-substituted hyaluronan macromolecules using a peroxidaseand H₂O₂ allows the resulting hydrogels to be cross-linked ex vivo, andto remain stable in animal tissues. For example, studies in rats havedemonstrated that this material does not degrade over several monthswhen injected subcutaneously (see Example 9). At low concentrations thehydrogels are optically clear, easily injected through a syringe or avitrectomy port and have a specific gravity higher than water. Thesephysical properties make T-HA gels an ideal substrate for vitreousreplacement.

Based in part on confined compression testing data reported in Example 3above, it was possible to design a T-HA hydrogel composed of, inter aliaa macromolecular network of dityramine cross-linked hyaluronanmolecules, having elastic and other physical properties matched tonatural vitreous material in order to produce a synthetic implantablevitreous substitute.

Choices of scaffold material (HA alone), percent tyramine substitution(˜5%), protocol for tyramine substitution of HA (modified from Example1), and exclusion of cells and bioactive factors were as described inExample 6. A concentration range of between 2.5 and 10 mg/ml of T-HAhydrogel in sterile saline most closely matched the rheologic, optical(clarity, refractive index) and gravimetric (density) properties of thevitreous of the eye. The 10 mg/ml concentration of T-HA hydrogel wasdeemed most appropriate based on the extensive clinical experience ofour clinician collaborators. In vitro cross-linked hydrogel was usedrather than an in situ cross-linking protocol based on the experience ofour clinician collaborators with the potential sensitivity of the thin,highly-specialized layer of cells in the retina. In vitro cross-linkingwas as described in Example 1. An insoluble steroid was added to thecross-linked T-HA to allow visualization by the surgeons during theoperative procedure because the hydrogel material itself was opticallytransparent and colorless.

Surgical Procedure

Rabbits underwent unilateral vitrectomy surgery (left eye only) usingstandard vitreoretinal surgical techniques with replacement of thenatural vitreous of the eye with the T-HA hydrogel described above inorder to evaluate the hydrogel material as a vitreous substitute.Following general anesthesia (ketamine: 50 mg/kg; xylazine: 5 mg/kg),the rabbit was prepped and draped in a sterile fashion. The left eye wasdilated with mydriacyl and phenylephrine. Two drops of topical Ciloxanwas instilled over the eye before and after the case. Topicalproparacaine drops were instilled. Under an operating microscope, a 270°conjunctival peritomy was performed using wescott scissors. An infusionport was created 2.5 mm posterior to the limbus and the infusion cannulawas secured to the sclera using 7-0 vicryl suture. A lens ring wassutured to the sclera using 7-0 vicryl suture. A 30° prism vitrectomylens was placed on the lens ring. A second port was created and thevitrectomy instrument was inserted into the vitreous cavity. A completecore and peripheral vitrectomy was performed. At this point one of theports was sutured closed with a 7-0 vicryl suture. The BSS bottle waslowered to patient level and ˜1.2 cc of a mixture of 3 mg/mlpreservative free triamcinolone acetonide (steroid) and the T-HAhydrogel (10 mg/ml, 5% tyramine substitution) in BSS was injected intothe vitreous cavity through an 18 gauge syringe. The steroid is the sameas usually administered after vitreoretinal surgery with its milkyappearance allowing visualization of the otherwise optically transparenthydrogel material. As the milky solution was injected it was directlyvisualized filling the vitreous cavity. The infusion was stopped when at50% fill or when the T-HA material was seen backing up through theirrigation canula (100% fill). After filling the vitreous cavity 7-0vicryl suture was used to close the remaining ports. 8-0 vicryl suturewas used to close the conjunctiva. Topical bacitracin ointment wasplaced on the eye after closing. A topical antibiotic/bacitracinointment was applied to the eye bid×1 week, and the rabbit was placed ina recovery cage. After the rabbit had regained sternal recumbancy, itwas returned to its home cage.

Post-Implantation Data

At 1 month post-implantation, rabbits were anesthetized with ketamine(50 mg/kg; 10 mg/kg/hr thereafter) and xylazine (5 mg/kg; 0.5 mg/kg/hrthereafter), the pupils dilated with eye drops (1% tropicamide; 2.5%phenylephrine), and the corneal surface anesthetized with an eye drop(0.5% proparacaine). After full pupil dilation, the status of the retinaand eye were examined by indirect opthalmoscopy followed by fundusphotography. In addition, intraocular pressure (IOP) was measured usinga Tonopen, a device that is used clinically and which makes minimalcontact with the corneal surface. Finally, the rabbit was placed on aheating pad in darkness for 1 hour, and electroretinograms (ERGs)recorded for both control and vitreous replaced eyes in response toflashes of light. ERG electrodes consist of a corneal contact lens andtwo platinum 0.5 inch Grass needle electrodes, placed in the cheek andtrunk. While still under anesthesia, the rabbit was euthanized by usingan intravenous dose of Beuthanasia D Special (1 ml/5 kg). Both controland vitreous replaced eyes were then enucleated and fixed in 10%buffered formalin for 24 hours, for histological evaluation.

The results at 1 month post-implantation, indicated minimalpost-operative inflammation of the surgically treated eye with normalIOP. By one week, a cataract had formed in the vitreous replaced eyerelative to the un-operated eye and to BSS control operated eyescreating a limited view of the posterior segment of the experimentaleye. Gross observation of the sectioned eyes showed the cataract in theanterior segment of the vitreous replaced eye (FIG. 10). The remainderof the anterior segment as well as the entire posterior segment of theeye looked similar by gross observation (FIG. 10). Hydrogel wasrecovered from the experimental eye at 1 month post-implantation, andwas a clear gel-like substance similar to its pre-injection form (FIG.10). The ERG for the vitreous replaced eye was normal compared to theun-operated control eye, and indicated that the retinal cells were aliveand remained functional (FIG. 11). Finally electron micrographs from thefour quadrants of the retina show normal morphology for the retina fromthe vitreous replaced eye compared to the normal un-operated eye (FIG.12). These data indicate the T-HA hydrogel can be used as a vitreoussubstitute without causing infection or inflammation of the eye, andwithout damaging the retina, and illustrate a method by which the T-HAhydrogel can be used as a retinal tamponade for reattachment of adetached retina.

Related Opthalmologic Applications

In addition to the foregoing retinal tamponade application followingretinal reattachment surgery, the T-HA hydrogel described in thisexample also could be used for the following related applications:

-   -   as a vitreous replacement for intra-operative procedures such as        unfolding of retinal tears, the removal of subretinal fluid and        floatation and removal of dislocated intraocular lens        components.    -   as a vitreous replacement incorporating a sustained release drug        delivery system to maintain therapeutic drug levels (steroids,        antibiotics, anti-viral drugs, etc.) in the posterior segment of        the eye over long periods of time to treat chronic inflammatory        and infectious conditions of the eye such as sarcoidosis,        idiopathic posterior uveitis and cytomegalovirus retinitis.    -   for anterior segment surgery including as a substitute for        plastic polymer inserts in corneal refractive surgery. These        inserts, implanted surgically in the cornea, are used to change        the shape of the cornea and correct mild myopia. The optical        clarity and biocompatibility of the T-HA hydrogel with human        tissues make it well suited to this application.    -   as a substitute for partial or full thickness corneal grafting        procedures, e.g. for anterior segment surgery, necessitated as a        result of corneal scarring from infection, keratinous, or other        causes. The optical and physical properties of the hydrogels        make them compatible with use as corneal tissue substitutes.    -   as a viscoelastic device during anterior segment and cataract        surgery. At low concentrations, hydrogels can maintain anterior        chamber shape and pressure while allowing the surgeon to clearly        visualize ocular structures.    -   for oculoplastic surgery including subcutaneous injection to        smooth wrinkles in the face.    -   for oculoplastic surgery as an ocular implant in patients        undergoing enucleation or exoneration surgery. The hydrogel        formed in the dimensions of a human eye can be used as an        implant to fill the orbit and improve cosmetic appearance of the        individual after globe removal.    -   to coat MEMS devices for use in vitreo-retinal surgery.    -   to expand the utility of laser vision correction surgery (LASIK)        to include those cases where volume needs to be added to the        cornea rather than removed to correct vision. Corrective laser        surgery would be used to produce the exact dimensions required        for optimal visual outcome following implantation of an        intentionally oversized plug of the hydrogel.    -   as a replacement for the typical gases, perfluorocarbon liquids        and silicon oils normally used as tamponades in eye surgery.        These applications include but are not limited to the following:        giant retinal tears, proliferative vitreoretinopathy (PVR),        large breaks with “fish-mouth” phenomenon, posterior breaks or        macular holes, the restoration of intraocular volume after        drainage of subretinal fluid, total retinal detachment with        multiple breaks and large meridional folds, retinal detachment        caused by ocular trauma or complicated by PVR or associated with        choroidal coloboma, dislocated lenses, suprachoroidal and        submacular hemorrhage, rhegmatogenous retinal detachments        without PVR, severe proliferative diabetic retinopathy, chronic        uveitis with profound hypotony, and infectious retinitis.

Example 9

An experiment was conducted whereby plugs of T-HA hydrogels as describedhereinabove were implanted subcutaneously into immunocompetent rats inorder to investigate and demonstrate their in vivo persistence andlongevity as well as to measure any host immune response. As describedin detail above, hydrogels comprising a cross-linked macromolecularnetwork (such as a tyramine-substituted and cross-linked hyaluronannetwork) can be prepared having a range of physical and viscoelasticproperties. These materials, for example, can be tuned to emulatenatural soft tissue and could be used for repair or augmentation of softtissue defects, as in plastic surgery or reconstructive surgery. Inparticular, as described in detail in Examples 3 and 4 above, theviscoelasticity, rigidity and other physical properties of the materialcan be tuned across a wide range to emulate like properties of a widevariety of native soft tissues, and the material can be cast or formedinto a variety of complex anatomical shapes which would make it idealfor casting replacement or reconstructed tissue components; e.g., in theshape of an ear or of a nose for facial reconstruction.

While it already was clear from the noted examples that these materialscould be cast into appropriate shapes and could be given appropriatephysical properties, the present experiment demonstrates the feasibilityof using the T-HA hydrogels as synthetic tissue matrix or replacementmaterials in vivo. Following is a description of the experiment,including the experimental methods and results obtained, after a briefdiscussion of the background for this application.

Background Tissue Description and the Need for a Synthetic Material

The availability of biomaterials for soft tissue augmentation and headand neck reconstruction has remained a fundamental challenge in thefield of plastic and reconstructive surgery. Significant research andinvestment has been undertaken for the development of a material withappropriate biological compatibility and life span. The present focus intissue engineering has been directed at attempts toward fibroblast andchondrocyte cultures as a method of creating endogenous cartilage andcollagen bearing structures useful for implantation. The archetypalstandard of this avenue of research has been the nude mouse with aneo-cartilage ear on its back. This is based on the concept ofchondrocyte culture on a poly-lactic or poly-glycolic acid framework.The presumption is that the chondrocytes can produce the extracellularmatrix (ECM) for the production of cartilage and create a new functionalbiological filling agent with complete compatibility. The outcomes ofthis research have not been promising in regards to their clinicalapplication. When placed in immunocompetent animals the structuralintegrity of the neo-cartilage has been shown to fail as the frameworkis absorbed. Fundamentally, while chondrocytes can successfully becultured and propagated they apparently cannot be made to producecartilage on a framework prior to its hydrolysis by the host defensemechanisms.

Conventionally, clinicians have been limited by the use of xenogenicmaterials such as bovine collagen and unmodified hyaluronan (HA) as wellas synthetic materials such as silicone, silastic and hydroxyapatite.The synthetic materials are prone to foreign body reactions andinfection while the biological substrates are prone to breakdown overtime. In addition, synthetic PTFE (gortex) polymers and silastic offerless tissue reactivity but do not offer tissue integration and also canrepresent long term risks of foreign body infections and extrusion.

Instead of a tissue engineering model where chondrocytes are required toproduce a cartilage ECM, the hydrogels disclosed herein are or can bebased on the same materials that provide cartilage its functionality andfeel (HA). In the present invention, hyaluronan is used directly as thesubstrate for the creation of a stable tissue engineered material toreplace natural HA-rich soft tissues. In essence, the HA-based hydrogelsused in herein incorporate the same material that gives cartilage itsform and structural characteristics, but it is modified(tyramine-substituted and cross-linked) to make the material resistantto biological degradation. Thus, an ideal synthetic extracellular matrixmaterial suitable for in vivo implantation and longevity is achieved.

Experimental Description

Design of Extracellular Matrix Material having Desired Properties

To provide a synthetic soft-tissue and cartilage substitute for use inhead and neck reconstruction based on the disclosed HA materials thefollowing points were considered: 1) optimization of enzyme-selectivecross-linked hydrogels using hyaluronan as the scaffold material; and 2)application of T-HA hydrogels as cartilage substitutes and soft-tissuefillers. These include characterization of the effect of thecross-linked hydrogels in vivo.

Again, based in part on confined compression testing data reported inExample 3 above, it was possible to design T-HA hydrogels composed of,inter alia a macromolecular network of dityramine cross-linkedhyaluronan molecules, having elastic and other physical propertiesmatched to natural soft tissues which were suitable for in, vivo ratimplantation to determine their immunogenic and longevitycharacteristics.

Choices of scaffold material (HA alone), percent tyramine substitution(˜5%), protocol for tyramine substitution of HA (modified from Example1), and exclusion of cells and bioactive factors were as described inExample 6. A concentration range of between 6.25 and 100 mg/ml of T-HAhydrogel encompassed the wide spectrum of physical properties requiredof a material for facial reconstruction. Therefore the same fiveconcentration used in Example 3 were deemed appropriate for testing in asubcutaneous rat model based on the extensive clinical experience of ourclinician collaborators. In vitro cross-linked hydrogels were used so asto produce hydrogels of defined shape for analysis of shape retention, aproperty deemed important by our clinician collaborators. In vitrocross-linking was as described in Example 1.

Surgical Procedure

T-HA hydrogel plugs of defined shape, mass and volume (7 mm in diameterand 3 mm in thickness) and defined mechanical properties based on HAconcentration were surgically implanted subcutaneously in the backs ofimmunocompetent rats to allow evaluation of their in vivo persistenceand host immune response based on previously published protocols for theevaluation of collagen and other HA-based hydrogels. After induction ofanesthesia with intraperitoneal injection of ketamine (100 mg/kg) andxylazine (5 mg/kg), the rat received a single intramuscular injection of60,000 units of procaine penicillin for infection prophylaxis. A 1 cmstab incision was made with a #11 surgical blade in the lower lumbarregion of the rat. A 14 g needle was used as a trocar to dissect in thesubcutaneous plane to create a pocket. Three preformed hydrogel plugs(˜7.1 mm diameter×3 mm thick) of one of the HA concentrations to betested were inserted into the surgical pocket. A single absorbablestitch (3-0 Chromic) was placed to re-approximate the skin edge. At 1week, 1 month, 3 months, and 6 months post-implantation, rats weresacrificed by CO₂ asphyxiation and the T-HA hydrogel plugs withsurrounding tissue excised and stored in formalin at 4° C. until timefor histological evaluation.

Post-Implantation Data

Hydrogel compositions tested included plugs made from concentrations of6.25, 12.5, 25, 50, and 100 mg/ml of HA, generating hydrogel plugs witha wide spectrum of physical properties ranging from that of gel to apaste to a rubber-like material (see Example 3). Implanted T-HA hydrogelplugs were collected at 1 week, 1 month, 3 months, and 6 monthspost-implantation. Excised plugs were evaluated for their in vivopersistence and host immune response.

FIG. 13 shows representative results of histological staining with H&E,alcian blue, MC/giemsa, Movats, Reticular, and Trichrome stains for the100 mg/ml TB-HA hydrogel plug from the 1-month time point. Clearlydefined in FIG. 13 are the surface hair follicles, the superficialmuscle layer, the hydrogel plug, and the thin fibrous capsulesurrounding the hydrogel plug as a result of a minimal foreign bodyresponse. An artifact exists as a result of hydrogel shrinkage from theparaffin embedding process, which can be avoided through the use offrozen sections. The results show very little immune response with onlya thin layer of mast cells surrounding the plug, and no evidence forhost cell infiltration into the plug. When measured, the volume of thevoid left by the plug during histological processing is 3 mm (theoriginal plug thickness) indicating little to no biodegradation ordeformation of the hydrogel matrix. Staining indicated that the plugshad little protein, such as collagen or elastin, deposited within themand remained primarily composed of HA hydrogel. These results indicatethat the hydrogel plugs over a broad range of five concentrationspersisted through 6 months with little evidence of degradation, hostimmune response, and cellular infiltration providing a wide range ofinjectable materials for use in soft tissue reconstruction.

Example 10

It will be apparent from the foregoing discussion and the Examples thathydrogels described herein composed of a cross-linked (in situ or exvivo) macromolecular network of hyaluronan molecules cross-linked via asuitable dihydroxyphenyl cross-linking chemistry as herein described aresuitable as a synthetic, implantable extracellular matrix tissuematerial for a variety of tissue engineering and repair applications. Aparticular such application for which the disclosed hydrogel materialswill have particular utility is in the repair or augmentation of themitral valve in a heart.

The mitral valve is one of the most complex connective tissue structuresin the body. It consists of two leaflets and numerous chordae tendineae.These chordae have a highly aligned collagenous core and a thin outersheath of elastic fibers and endothelial cells. Both leaflets arelaminated tissues containing a heavily collagenous layer on theventricular side, a predominantly elastic layer on the atrial side, andan inner spongiosa layer containing abundant proteoglycans (PGs) andhyaluronan (HA). The relative thicknesses of these layers vary betweenthe two leaflets and also within each leaflet from its attachment edgeto its free edge.

The variability of the different leaflet layers, and hence thestructural constituents within the mitral valve, are determined by thespecific functional roles of the leaflets and chordae. The closed valvemaintains a balance of tensile and compressive loads, in which thechordae and the flat central region of the anterior leaflet are intension, whereas the free edge of the anterior leaflet and most of theposterior leaflet are in appositional compression. Accordingly, the mostcollagenous components of the mitral apparatus are the chordae and theportion of the anterior leaflet between the annulus and the upperappositional border. In the posterior leaflet and in the free edge ofthe anterior leaflet, the collagenenous layer is relatively thinner,whereas the PG rich spongiosa is substantially thicker. The widediversity of glycosaminoglycans (GAGs) and their parent PGs exertconsiderable yet variable control over the physical properties of theextracellular matrix.

Functional mitral regurgitation (MR) refers to the regurgitation thatoccurs with a structurally normal valve as a consequence of leftventricular (LV) dysfunction, and as a result, almost half the patientswith LV dysfunction have at least moderate MR. Functional MR plays apivotal role in the pathophysiology of congestive heart failure (CHF), amajor cause of cardiac morbidity and mortality. Several studies haveshown that the presence of functional MR in patients with CHF isassociated with poor outcomes. Although this observation could suggestthat MR is merely an indicator of CHF severity, it is also increasinglyapparent that the development of the MR hastens the progression of CHF.The precise mechanism of functional MR remains controversial and canrelate to mitral annular dilatation in the septal-lateral (S-L) axis ortethering of the leaflets secondary to progressive ventricularremodeling. MR leads to greater volume overload of the LV withprogressive annular dilatation and increased MR, creating a “viciouscycle” which exacerbates the problem. MR is commonly considered to beone of the initiators of CHF, as well as an ongoing impetus of theprogression of the disease.

Surgical annuloplasty is a widely used method for mitral valve repairand can provide long-term benefit. However, the surgical procedurerequires access to and manipulation of the valve annulus via atriotomy.In addition, the procedure requires the patient to be placed oncardiopulmonary bypass (CPB). The prolonged CPB time has been suggestedas a cause of not only postoperative LV dysfunction but also main organdysfunction. The use of heparin during CPB results in an increased riskof bleeding complications. The increased morbidity and mortality profileleads many care providers directly to non-treatment options of MR in theearlier stage heart failure patients.

Recently, several minimally invasive methods of mitral valve repair havebeen developed. For example, several investigators have reported thepreliminary methodology of off-pump mitral valve repair proceduresthrough a thoracic incision. Others have reported new devices that canbe inserted percutaneously into the coronary sinus and great cardiacvein to reduce the S-L dimension of the mitral annulus. There are thepossibilities of adverse effects, however, such as obstruction ordisturbance of the coronary circulation, by chronic placement of thedevice in the coronary sinus.

Surgical therapy for functional MR, including mitral valve repair withan annuloplasty ring and replacement with an artificial valve has beenlimited in patients with severe CHF by relatively high operativemortality rates due to the effects of CPB. Therefore, there is a needfor a minimally invasive procedure that will not compromise coronarycirculation and will allow for reduction of the S-L dimension of themitral annulus to reduce functional MR as well as other forms of MR.Myxomatous changes in mitral valve tissues can lead to leaflet prolapseand mitral regurgitation.

The T-HA hydrogel materials disclosed herein could be adapted for thispurpose; i.e. mitral annular remodeling resulting from a nonabsorbablesubstance injection (namely a T-HA hydrogel material designed to havethe necessary viscoelastic and other physical properties) into theposterior mitral annulus using an epicardial approach. This procedurewould enable the S-L dimension to be efficiently reduced, thus reducingMR, without employing CPB or implanting a device into the coronarysinus. This mitral annular remodeling procedure could be modified toallow for percutaneously injection of the substance through the coronarysinus.

This application for a T-HA hydrogel as disclosed herein would enable anonabsorbable substance to be percutaneously injected through thecoronary sinus for severe CHF patients with functional MR who are unableto receive conventional mitral valve surgery. This minimally invasiveapproach would obviate the need for CPB and sternotomy, as well asdiminish the risk of major side effects from conventional surgicaltherapy such as postoperative LV dysfunction and resulting poor organperfusion. In addition, such a procedure would provide patients withmild to moderate CHF with an option for early restoration of mitralvalve competence, arresting the initiation and progression ofdevastating heart failure.

In particular, a hydrogel composed of a cross-linked macromolecularnetwork as described herein, particularly of HA, could be designed basedon the principles as elucidated in Example 3 above, in order to producea hydrogel material having all of the following characteristics whichwould be considered desirable for this application:

Injectable and nonabsorbable;

Low-grade inflammatory reaction;

Low evidence of foreign body migration;

Ease of collagen encapsulation which contributes to the prevention ofmigration;

Not especially malleable nor especially rigid.

A protocol has been established for the injection of a T-HA hydrogelmaterial to augment the mitral valve of a beating heart. That protocolis described as follows.

Injection Procedure—Two dimensional epicardial echocardiography (2D EE)and transesophageal echocardiography (2D TEE) are performed to evaluateLV end-diastolic and end-systolic volumes (EDV and ESV), stroke volume(SV), ejection fraction (EF), the S-L dimension of mitral annulus, andthe degree of MR. Hemodynamic data such as LVP, LAP, the central venouspressure (CVP), the pulmonary arterial pressure (PAP), the pulmonarycapillary wedge pressure (PCWP), CO, LAD flow, and LCX flow should becollected. LV end-diastolic and end-systolic pressure-volume relationscan be obtained by transient IVC occlusion using an occlusion catheterto assess LV contractility and compliance (baseline before injection).

A commercially available cardiac stabilizer used for off-pump coronarybypass grafting can be used to stabilize the target region. Under 2D TEEguidance, the uncross-linked T-HA hydrogel composition (50 mg/ml insterile saline) is injected into the posterior mitral annulus from theoutside of the heart while the heart is beating. During the injection,2D can be used to assess the position of the tip of the needle for theinjection, the range occupied by the substance in the posterior annulus.Once an appropriate fill and repositioning of the mitral valve has beenaccomplished, cross-linking is initiated by injection of 0.2 cc of 0.6%hydrogen peroxide. After completion of cross-linking of the hydrogelimplant, data including hemodynamics, coronary flow, LV pressure-volumeloops (LV P-V loops), 2D EE, and 2D TEE should be collected (data afterinjection).

The foregoing injection methodology was developed and evaluated usingcadaveric dog and pig hearts as models. FIG. 14 shows a cadaveric dogheart in which a T-HA hydrogel material was injected and cross-linked insitu via the foregoing injection methodology. To produce the T-HAhydrogel material for this experiment, the scaffold material (HA alone),percent tyramine substitution (˜5%), protocol for tyramine substitutionof HA (modified from Example 1) and non-inclusion of cells and bioactivefactors were as described in Example 6. Concentrations of 25, 50 and 100mg/ml of T-HA hydrogel in saline most closely matched those of cardiac(heart) tissue required for mitral valve closure. The 50 mg/ml T-HAhydrogel was deemed most appropriate for a mitral annular remodelingprocedure by percutaneously injection based on the extensive clinicalexperience of our clinician collaborators. Peroxidase would be added at10 U/ml prior to application in an in situ cross-linking protocol as theone described below. The in situ cross-linking protocol is preferred asit would allow the uncross-linked T-HA to pass through an appropriatelysized needle while the cross-linked hydrogel may not. In addition, thein situ cross-linking protocol allows the surgeon first to properlyposition (close) the mitral valves by injection of the uncross-linkedhydrogel and then cross-link the hydrogel into a solid implant onlyafter visual confirmation that the mitral valve had been properlyrepositioned.

In FIG. 14, the implanted hydrogel was bisected post-implantation todemonstrate its solid, viscoelastic character following in situcross-linking, as well as to evaluate its placement in the heart. Inparticular, these models were used to 1) evaluate the appropriateconcentration of hydrogel required to both mimic the consistency ofcardiac muscle after cross-linking yet pass through the injection portprior to cross-linking; 2) demonstrate reproducibility for the in vivocross-linking protocols for complete cross-linking in vivo at therequired volumes (˜2 ml); and development suitable injection techniques.All of these goals were met, establishing confidence that the injectionprocedure as well as a suitable T-HA hydrogel can precisely accommodateanatomical constraint on the mitral annulus.

Although the above-described embodiments constitute the preferredembodiments, it will be understood that various changes or modificationscan be made thereto without departing from the spirit and the scope ofthe present invention as set forth in the appended claims.

1. A method of making a hydrogel in situ comprising the steps of: a)providing a first solution comprising either a peroxidase enzyme or aperoxide but not both, and hydroxyphenyl-substituted hyaluronanmolecules; b) providing a second solution comprising either theperoxidase enzyme or peroxide not provided in the first solution; and c)combining the first and second solutions in situ to initiatedihydroxyphenyl cross-linking between the hydroxyphenyl-substitutedhyaluronan molecules to form the hydrogel.
 2. The method of claim 1,further comprising the step of integrating the first or second solution,or both, into tissue at a location of interest in situ within a personor animal.
 3. The method of claim 1, wherein the concentration ofhydroxyphenyl-substituted hyaluronan molecules in the first and secondsolutions, when combined, is between 6.25 and 100 mg per mL.
 4. Themethod of claim 1, wherein the concentration ofhydroxyphenyl-substituted hyaluronan molecules in the first and secondsolutions, when combined, is between 6.25 and 25 mg per mL, and therigidity, rheology and texture of the hydrogel ranges from that of ajelly-like composition to a dough-like composition.
 5. The method ofclaim 1, wherein the concentration of hydroxyphenyl-substitutedhyaluronan molecules in the first and second solutions, when combined,is 25 and 100 mg per mL, and the rigidity, rheology and texture of thehydrogel ranges from that of a dough-like composition to acartilage-like material.
 6. The method of claim 1, wherein thehydroxyphenyl-substituted hyaluronan molecules are tyrosine-substitutedhyaluronan molecules.
 7. The method of claim 1, wherein one or both ofthe first and second solutions includes a population of viable livingcells.
 8. The method of claim 1, wherein the hydrogel that is formedincludes a population of viable living cells.
 9. The method of claim 8,wherein the population of viable living cells include cells selectedfrom the group consisting of chondrocytes, progenitor cells, and stemcells.
 10. The macromolecular network of claim 1, wherein one or both ofthe first and second solutions includes one or more bioactive factors.11. The method of claim 10, wherein the bioactive factors includebioactive factors selected from the group consisting of growth factors,hormones, and factors controlling cell differentiation.
 12. The methodof claim 1, wherein the in situ hydrogel is formed in an animal and thehydrogel exhibits little or no degradation after one month within theanimal.
 13. A method of making a hydrogel in situ comprising the stepsof: a) providing a first solution comprising either a peroxidase enzymeor a peroxide but not both, and tyrosine-substituted hyaluronanmolecules; b) providing a second solution comprising either theperoxidase enzyme or peroxide not provided in the first solution; c)integrating the first or second solution, or both, into tissue at alocation of interest in situ within a person or animal; and d) combiningthe first and second solutions in situ to initiate dityrosinecross-linking between the tyrosine-substituted hyaluronan molecules toform the hydrogel; wherein the concentration of tyrosine-substitutedhyaluronan molecules in the first and second solutions, when combined,is between 6.25 and 100 mg per mL, wherein the resulting hydrogelincludes the population of viable living cells.
 14. The method of claim13, wherein the concentration of tyrosine-substituted hyaluronanmolecules in the first and second solutions, when combined, is between6.25 and 25 mg per mL, and the rigidity, rheology and texture of thehydrogel range from that of a jelly-like composition to a dough-likecomposition.
 15. The method of claim 13, wherein the concentration oftyrosine -substituted hyaluronan molecules in the first and secondsolutions, when combined, is between 25 and 100 mg per mL, and therigidity, rheology and texture of the hydrogel range from that of adough-like composition to a cartilage-like material.
 16. The method ofclaim 13, wherein the population of viable living cells includes cellsselected from the group consisting of chondrocytes, progenitor cells,and stem cells.
 17. The method of claim 13, wherein one or both of thefirst and second solutions includes one or more bioactive factors. 18.The method of claim 17, wherein the bioactive factors include bioactivefactors selected from the group consisting of growth factors, hormones,and factors controlling cell differentiation.
 19. The method of claim13, wherein one or both of the first and second solutions includes atleast a portion of the population of viable living cells.
 20. The methodof claim 13, wherein the in situ hydrogel that is formed exhibits littleor no degradation after one month within the person or animal.